Non-invasive and wearable chemical sensors and biosensors

ABSTRACT

A non-invasive epidermal electrochemical sensor device includes an adhesive membrane; a flexible or stretchable substrate disposed over the adhesive membrane; and an anodic electrode assembly disposed over the flexible or stretchable substrate including an iontophoretic electrode. The device includes a cathodic electrode assembly disposed adjacent to the anodic electrode assembly over the flexible or stretchable substrate and includes an iontophoretic electrode. Either the cathodic electrode assembly or the anodic electrode assembly also includes a sensing electrode that includes a working electrode and at least one of a counter electrode or a reference electrode. The iontophoretic electrode in either the anodic electrode assembly or the cathodic electrode assembly that includes the sensing electrode is disposed on the substrate to at least partially encompass the working electrode and the at least one of the counter electrode or the reference electrode. The device includes an electrode interface assembly including independent electrically conductive contacts.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of and claims priority to U.S. patentapplication Ser. No. 15/531,982, filed on May 31, 2017, which is a § 371National Stage application of International Application No.PCT/US2015/063836, entitled “NON-INVASIVE AND WEARABLE CHEMICAL SENSORSAND BIOSENSORS,” filed on Dec. 3, 2015, which claims priorities to andbenefits of U.S. Provisional Patent Application No. 62/087,172 entitled“NON-INVASIVE AND WEARABLE GLUCOSE MONITORING SENSORS” filed on Dec. 3,2014, and U.S. Provisional Patent Application No. 62/112,608 entitled“SINGLE-USE NON-INVASIVE WEARABLE ELECTROCHEMICAL SENSORS BASED ON“PLACE-DETECT-DISPOSE” (“PDD”) OPERATION “filed on Feb. 5, 2015. Theentire content of the above patent applications is incorporated byreference as part of the disclosure of this patent document.

TECHNICAL FIELD

This patent document relates to electrochemical sensor devices, systems,and techniques.

BACKGROUND

Various medical conditions of patients need regular or frequentpatient-operated testing to monitor conditions or well-being ofpatients. One example for such patient-operated testing is measuring theglucose level of diabetes patients. Diabetes (diabetes mellitus) is ametabolic disease associated with high blood sugar due to insufficientproduction of insulin by the body or inadequate response by cells to theinsulin that is produced. There are three types of diabetes: type 1,type 2, and gestational diabetes. Type 1 diabetes is associated with thebody's failure to produce insulin. Type 2 diabetes is associated withinsulin resistance, in which cells fail to use insulin properly. Thethird form of diabetes is referred to as gestational diabetes, which canoccur when pregnant women develop a high blood glucose level (e.g., evenwithout a previous history of diabetes). Gestational diabetes candevelop into type 2 diabetes, but often resolves after the pregnancy.Diabetes is widely-spread globally, affecting hundreds of millions ofpeople, and is among the leading causes of deaths globally. Variousmedical conditions of patients need regular or frequent patient-operatedtesting to monitor conditions or well-being of patients. One example forsuch patient-operated testing is measuring the glucose level of adiabetes patients.

SUMMARY

Techniques, systems, and devices are disclosed for non-invasive andwearable biosensing or chemical monitoring, such as glucose monitoring.The disclosed technology enables monitoring in continuous orquasi-continuous mode through multiple measurements performed using thesame sensor over a time period (longer than measured in minutes, e.g., afew hours to several days). A continuous chemical monitoring device canbe obtained by selecting a hydrogel layer that allows fast transport ofall the extracted chemical analyte to the sensor surface. The basicimplementation of the disclosed technology is a single-use on-bodyelectrochemical sensors for non-invasive detection of physiologicallyrelevant chemicals. For example, the disclosed technology provides aneconomical and easy-to-use biosensor platform that allows a user to“place-detect-dispose” (PDD)’ the sensor to obtain bioanalyticalinformation for health and performance monitoring. However, thedisclosed biosensor is reusable for multiple testing that allows for thecontinuous or quasi-continuous monitoring of a target analyte. Thus, thePPD application is described in this patent document as an exemplary usecase and a variety of reusable protocols can be implemented toperforming a target analyte monitoring over a period of timecontinuously or periodically. The disclosed technology can detectpre-selected chemicals in the skin interstitial fluid (ISF) or in sweatat a particular time to allow for non-invasive and easy measurementsthat can be carried out by the patients with ease without cutting orpiercing skin or any body tissues of the patients. For example, thedisclosed technology can be implemented as disposable, non-invasiveelectrochemical sensors to be conveniently carried by the patents attimes of their choice, e.g., after a meal, such as for glucose testing,or after beverage consumption, such as for alcohol testing. For example,a user can apply an exemplary single-use electrochemical sensor deviceto a selected skin area such that the device then quantifies a targetchemical analyte in a biological fluid, e.g., including an analyte inISF that is non-invasively extracted from under the epidermal layer ofthe skin for external detection by the device. Thereafter, the user canremove the device from the skin.

The subject matter described in this patent document can be implementedin specific ways that provide one or more of the following features. Thedisclosed technology can be used for a variety of applications includingdigital/mobile health applications; its customizable consumable lendsitself to a range of business models and services. For example, inapplications of detecting chemicals in the ISF, an electrochemicalsensor device includes a reverse iontophoretic electrode assembly thatwill first extract the ISF to the surface of the skin followed byelectrochemical detection of the desired chemical analyte viaelectrochemical sensor electrodes of the device. In applications ofdetecting chemicals in the sweat, an iontophoretic electrode assembly isincluded into the device that will first induce localized sweating (inthe vicinity of the device) by transdermally administeringsweat-inducing drugs into the skin, followed by detection of thechemical analyte in the induced sweat via the electrochemical sensorelectrodes. Also, for example, the device can be configured with visualdesigns that can be customized for the purpose of branding, aesthetics,user specification (user's own design), time/condition labeling etc.

In one aspect, a non-invasive epidermal electrochemical sensor deviceincludes a flexible substrate including an electrically insulativematerial structured to adhere to skin of a user; an anodic electrodeassembly and a cathodic electrode assembly each formed and separatelyarranged on the substrate and each including a working electrode, acounter/reference electrode, and a reverse iontophoretic electrode, inwhich the reverse iontophoretic electrode is structured on the substrateto at least partially encompass the working electrode and thecounter/reference electrode and operable to apply an electric field todrive ion flow from interstitial fluid (ISF) toward the workingelectrode and the counter/reference electrode, and in which the workingelectrode includes a electrochemical transducer layer including acatalyst to selectively catalyze a corresponding analyte in the ISF tocause a reaction detectable at the anodic and cathodic electrodeassemblies; and an electrode interface assembly comprising independentelectrically conductive conduits formed on the substrate andelectrically coupled to each of the working, counter/reference, andreverse iontophoretic electrodes of the anodic electrode assembly andthe cathodic electrode assembly. The non-invasive epidermalelectrochemical sensor device, when attached to the skin andelectrically coupled to one or more electrical circuits via theelectrode interface assembly, the device is operable to detect theanalyte from the ISF in a local environment of the skin.

The subject matter described in this patent document can be implementedin specific ways that provide one or more of the following features. Insome aspects, a glucose monitoring device of the disclosed technologyincludes an all-printed temporary tattoo-based glucose sensor fornon-invasive glycemic monitoring. The exemplary flexible tattoo-basedepidermal diagnostic device uses reverse iontophoretic extraction ofinterstitial glucose and an enzyme-based amperometric biosensor.Exemplary in vitro implementations of the flexible tattoo-basedepidermal glucose monitoring device produced exemplary data showing thetattoo sensor's linear response towards physiologically relevant glucoselevels with negligible interferences from common co-existingelectroactive species. The disclosed iontophoretic-biosensing tattoosensor technology platform can be applied on human subjects, e.g., byattaching to the subject's skin, to monitor variations in glycemiclevels due to food consumption. Also, for example, exemplaryimplementations to compare performance of the disclosed non-invasive andwearable glucose monitoring technology with conventional glucose meterswere performed, in which the exemplary results included correlation ofthe disclosed sensor response with that of a commercial glucose meter,which underscores the benefits of the exemplary tattoo sensor to detectglucose levels in a non-invasive fashion. The exemplary implementationsincluded control on-body experiments that demonstrated the importance ofthe reverse iontophoresis operation and validate the sensor specificity.The disclosed tattoo-based iontophoresis-sensor platform holdsconsiderable promise for efficient diabetes management and can beextended towards non-invasive monitoring of other physiologicallyrelevant analytes present in the interstitial fluid.

In one aspect, a non-invasive epidermal biosensor device includes aflexible substrate including an electrically insulative materialstructured to adhere to skin of a user; an anodic electrode assembly anda cathodic electrode assembly disposed on the flexible substrate, inwhich each of the anodic and cathodic electrode assemblies includes (i)electrochemical sensing electrodes including a working electrode and asecond electrode and being configured to form an electrochemical sensor,and (ii) an iontophoretic electrode disposed on the flexible substrateto at least partially encompass an area occupied by the workingelectrode to apply an electric field to a skin area of the userincluding a tissue area underneath the skin surface to drive ions flowin interstitial fluid (ISF) underneath the skin surface toward theelectrochemical sensing electrodes, in which the iontophoreticelectrodes are configured not part of the electrochemical sensingelectrodes, and in which the working electrode of the cathode and/or theanode electrode assembly includes a electrochemical transducer layerincluding a catalyst to selectively catalyze a corresponding analyte tocause a reaction detectable at the electrochemical sensing electrodes ofthe anodic and cathodic electrode assemblies; a layer disposed over theiontophoretic electrodes of the anodic and cathodic electrode assembliesto provide an electrically conductive medium between the iontophoreticelectrodes and the skin for the electric field applied; and an electrodeinterface assembly comprising independent electrically conductiveconduits disposed on the flexible substrate that are electricallycoupled to the anodic and cathodic electrode assemblies to electricallycouple the anodic and cathodic electrode assemblies to one or moreexternal electrical circuits to electrically energize theelectrochemical sensing electrodes and to separately electricallyenergize the iontophoretic electrode. When attached to the skin, thedevice is operable to detect the analyte from the ISF in a localenvironment of the skin.

Implementations of the device can include one or more of the followingfeatures. In some implementations, for example, the analyte is glucose,and the catalyst can include glucose oxidase (GOx) or glucosedehydrogenase (GDH). For example, the electrochemical transducer layercan include a solution of GOx or GDH and a solution of a biopolymer, forexample, chitosan to immobilize the GOx or GDH in the electrochemicaltransducer layer. For example, the device is operable to extract theglucose from the ISF in a region containing the electrochemical sensingelectrodes at the cathodic electrode assembly to cause the workingelectrode of the cathodic electrode assembly to react with (e.g.,electrochemically oxidize) the glucose via the GOx or GDH for selectiveglucose detection. In some implementations, for example, the flexiblesubstrate includes paper. In some implementations, for example, theflexible substrate includes plastic. In some implementations, forexample, the substrate includes fabric. In some implementations, forexample, the flexible or stretchable substrate includes at least one ofsilicone or polyurethane. In some implementations, for example, thesecond electrode of the electrochemical sensing electrodes and theiontophoretic electrode are structured as Ag/AgCl electrodes. In someimplementations, for example, the electrochemical sensor includes athird electrochemical sensing electrode, in which the third electrodecan be of same material or of different material as compared to theworking electrode. For example, in a two electrode configuration of theelectrochemical sensor, the second electrode is operable as a referenceand counter electrode. Also, for example, in a three electrodeconfiguration of the electrochemical sensor, the second electrode isoperable as a reference electrode, and the third electrode is operableas a counter electrode. In some implementations, for example, theworking electrode includes a hydrogen peroxide sensing transducer, suchas Prussian Blue. In some implementations, for example, the workingelectrode includes metal. In some implementations, for example, theworking electrode includes carbon. In some implementations, for example,the working electrode includes mediators for efficient electron transferbetween enzyme and electrode. In some implementations, for example, theiontophoretic electrode is arranged to extract interstitial fluid (ISF)in a region within 10 mm or less of the working electrode. In someimplementations, for example, the device further includes a firstinsulating layer formed over the anodic electrode assembly and a secondinsulating layer formed over the cathodic electrode assembly, the firstand second insulating layers to confine electrode and contact areas ofthe device. In some implementations, for example, the layer includes ahydrogel layer. For example, the layer (e.g., the hydrogel layer) canalso be coated over at least some of the electrochemical sensingelectrodes (e.g., the working electrode) of the anodic and cathodicelectrode assemblies to assist the flow of the extracted ISF to theelectrochemical sensor.

In one aspect, a method of non-invasively detecting an analyte using annon-invasive electrochemical sensor includes attaching anelectrochemical sensor device to a user's skin, in which theelectrochemical sensor device includes: (i) a flexible or stretchableand electrically insulative substrate, (ii) an anode electrode assemblyand a cathode electrode assembly separately disposed on the substrateand each including an iontophoretic electrode and two or moreelectrochemical sensor electrodes, which include a working electrode, inwhich the iontophoretic electrode is structured on the substrate to atleast partially encompass the working electrode, and in which theworking electrode of the cathode and/or anode electrode assemblyincludes a electrochemical transducer layer including a catalyst toselectively catalyze a corresponding analyte to cause a reactiondetectable at the electrochemical sensor electrodes of the anode andcathode electrode assemblies, and (iii) hydrogel layer formed over theiontophoretic electrodes of the anode and cathode electrode assembliesto provide an electrically conductive medium between the iontophoreticelectrodes and the skin. The attaching includes placing the anode andthe cathode electrode assemblies in contact with the skin, in which thesubstrate of the electrochemical sensor device adheres to the skin. Themethod includes extracting interstitial fluid (ISF) containing theanalyte onto the skin by applying an electric field from theiontophoretic electrodes of the anode and the cathode electrodeassemblies to drive ion flow of the ISF under an epidermal layer of theskin toward the electrochemical sensor electrodes. The method includesproviding an electrical potential at the working electrode to cause aredox reaction with the analyte in the extracted ISF that produces anelectrical signal. And, the method includes detecting the electricalsignal at the electrochemical sensor electrodes, in which the electricalsignal is associated with a parameter of the analyte, and in which thedetecting includes the oxidation or reduction of the analyte by thecatalyst (for example, enzyme) to generate electrical signal.

Implementations of the method can include one or more of the followingfeatures. In some implementations, for example, the analyte is glucose,and the catalyst can include glucose oxidase (GOx) or glucosedehydrogenase (GDH). In some implementations, for example, the methodfurther includes processing the detected electrical signal to determinea concentration level of the glucose in the ISF.

In one aspect, a non-invasive epidermal biosensor device to detect ananalyte in a biological fluid includes a flexible or stretchablesubstrate including an electrically insulative material structured toadhere to skin of a user; an anodic electrode assembly and a cathodicelectrode assembly disposed on the flexible substrate, in which each ofthe anodic and cathodic electrode assemblies includes (i)electrochemical sensing electrodes including a working electrode and asecond electrode being configured to form an electrochemical sensor, and(ii) an iontophoretic electrode disposed on the flexible substrate to atleast partially encompass an area occupied by the working electrode toapply an electric field into the skin, in which the iontophoreticelectrodes are configured not part of the electrochemical sensingelectrodes, and in which the working electrode of the cathode and/oranode electrode assembly includes a electrochemical transducer layerincluding a chemical agent to selectively catalyze or react with acorresponding analyte to cause a reaction detectable at theelectrochemical sensing electrodes of the anodic and cathodic electrodeassemblies; a layer disposed over the iontophoretic electrodes of theanodic and cathodic electrode assemblies to provide an electricallyconductive medium between the iontophoretic electrodes and the skin forthe electric field applied, in which the layer of the anodic electrodeassembly contains a chemical compound capable of inducing perspirationof the skin, such that, when the electric field is applied by theiontophoretic electrodes, the electric field causes the release of thechemical compound from the layer into the skin; and an electrodeinterface assembly comprising independent electrically conductiveconduits disposed on the flexible substrate that are electricallycoupled to the anodic and cathodic electrode assemblies to electricallycouple the anodic and cathodic electrode assemblies to one or moreexternal electrical circuits to electrically energize theelectrochemical sensing electrodes and to separately electricallyenergize the iontophoretic electrode. When attached to the skin, thedevice is operable to detect the analyte in sweat in a local environmentof the skin that was induced by the device.

Implementations of the device can include one or more of the followingfeatures. In some implementations, for example, the chemical compoundcan include pilocarpine. In some implementations, for example, theanalyte includes alcohol or an electrolyte. In some implementations, forexample, the chemical agent includes sodium nitrate. In someimplementations, for example, the flexible substrate includes paper. Insome implementations, for example, the substrate includes fabric. Insome implementations, for example, the flexible or stretchable substrateincludes at least one of silicone or polyurethane. In someimplementations, for example, the flexible substrate includes plastic.In some implementations, for example, the second electrode of theelectrochemical sensing electrodes and the iontophoretic electrode arestructured as Ag/AgCl electrodes. In some implementations, for example,the electrochemical sensor includes a third electrochemical sensingelectrode, in which the third electrode can be of same material or ofdifferent material as compared to the working electrode. For example, ina two electrode configuration of the electrochemical sensor, the secondelectrode is operable as a reference and counter electrode. Also, forexample, in a three electrode configuration of the electrochemicalsensor, the second electrode is operable as a reference electrode, andthe third electrode is operable as a counter electrode. In someimplementations, for example, the working electrode includes a hydrogenperoxide sensing transducer, such as Prussian Blue. In someimplementations, for example, the device further includes a firstinsulating layer formed over the anodic electrode assembly and a secondinsulating layer formed over the cathodic electrode assembly, the firstand second insulating layers to confine electrode and contact areas ofthe device. In some implementations, for example, the layer includes ahydrogel layer. For example, the layer (e.g., the hydrogel layer) canalso be disposed over the at least some of the electrochemical sensingelectrodes (e.g., the working electrode) of the anodic and cathodicelectrode assemblies to assist the flow of the extracted ISF to theelectrochemical sensor.

In one aspect, a method of non-invasively detecting an analyte in abiological fluid using an non-invasive electrochemical sensor includesattaching an electrochemical sensor device to a user's skin, in whichelectrochemical sensor device includes: (i) a flexible or stretchableand electrically insulative substrate, (ii) an anode electrode assemblyand a cathode electrode assembly separately disposed on the substrateand each including an iontophoretic electrode and two or moreelectrochemical sensor electrodes, which include a working electrode, inwhich the iontophoretic electrode is structured on the substrate to atleast partially encompass the working electrode, and in which theworking electrode of the cathode and/or anode electrode assemblyincludes a electrochemical transducer layer including a chemical agentto selectively catalyze or react with a corresponding analyte to cause areaction detectable at the electrochemical sensor electrodes of theanode and cathode electrode assemblies, and (iii) hydrogel layer formedover the iontophoretic electrodes of the anode and cathode electrodeassemblies to provide an electrically conductive medium between theiontophoretic electrodes and the skin. The attaching includes placingthe anode and the cathode electrode assemblies in contact with the skin,in which the substrate of the electrochemical sensor device adheres tothe skin. The method includes applying an electric field from theiontophoretic electrodes of the anode and the cathode electrodeassemblies to cause the release of the chemical compound from thehydrogel layer into the skin to induce perspiration onto the skincontaining the analyte. The method includes providing an electricalpotential at the working electrode to cause a redox reaction with theanalyte in the perspiration on the skin that produces an electricalsignal. And, the method includes detecting the electrical signal at theelectrochemical sensor electrodes, in which the electrical signal isassociated with a parameter of the analyte.

Implementations of the method can include one or more of the followingfeatures. In some implementations, for example, the detecting includesreleasing the catalyst corresponding to the analyte into the localenvironment to catalytically or meditatively participate in the redoxreaction. In some implementations, for example, the chemical compoundincludes pilocarpine, and the analyte includes alcohol or anelectrolyte. In some implementations, for example, the method furtherincludes processing the detected electrical signal to determine aconcentration level of the alcohol or the electrolyte in theperspiration.

In one aspect, a non-invasive epidermal biosensor device includes aflexible substrate to adhere to skin of a user; an iontophoreticelectrode assembly on the flexible substrate to apply an electric fieldto a skin area of the user including a tissue area underneath the skinsurface to extract interstitial fluid (ISF) from underneath the skinsurface outward from the skin; an electrochemical sensor comprising ananodic electrode assembly and a cathodic electrode assembly on theflexible substrate to sense an electrical signal from an electrochemicalreaction involving an analyte in the extracted ISF; a layer disposedover the iontophoretic electrodes to provide an electrically conductivemedium between the iontophoretic electrodes and the skin for theelectric field applied; and an electrode interface assembly comprisingindependent electrically conductive conduits on the flexible substratethat electrically couple the anodic and cathodic electrode assemblies ofthe electrochemical sensor and the iontophoretic electrode assembly toone or more external electrical circuits, respectively, to electricallyenergize the electrochemical sensing electrodes and to separatelyelectrically energize the iontophoretic electrode. When attached to theskin, the device is operable to detect the analyte from the ISF in alocal environment of the skin.

Implementations of the device can include one or more of the followingfeatures. In some implementations, for example, each of the anodic andcathodic electrode assemblies includes electrochemical sensingelectrodes including a working electrode and a second electrode. In someimplementations, for example, the electrochemical sensor electrodesinclude a third electrode, in which the third electrode can be of samematerial or of different material as compared to the working electrode.For example, in a two electrode configuration of the electrochemicalsensing electrodes, the second electrode is operable as a reference andcounter electrode. Also, for example, in a three electrode configurationof the electrochemical sensing electrodes, the second electrode isoperable as a reference electrode, and the third electrode is operableas a counter electrode. In some implementations, for example, theiontophoretic electrode assembly is disposed on the flexible substrateto at least partially encompass an area occupied by the workingelectrode. In some implementations, for example, the working electrodeof the cathode and/or anode electrode assembly includes aelectrochemical transducer layer including a catalyst to selectivelycatalyze a corresponding analyte to cause a reaction detectable at theelectrochemical sensor. In some implementations, for example, theanalyte includes glucose, and the catalyst includes glucose oxidase(GOx). In some implementations, for example, the electrochemicaltransducer layer includes a solution of GOx and a solution of chitosanto immobilize the GOx in the electrochemical transducer layer. In someimplementations, for example, the device is operable to extract theglucose from the ISF in a region containing the electrochemical sensingelectrodes at the cathodic electrode assembly to cause the workingelectrode of the cathodic electrode assembly to modify (e.g.,electrochemically oxidize) the glucose via the GOx for selective glucosedetection. In implementations, for example, the iontophoretic electrodeassembly is configured to not be part of the electrochemical sensor. Insome implementations, for example, the device further includes a firstinsulating layer formed over the anodic electrode assembly and a secondinsulating layer formed over the cathodic electrode assembly, the firstand second insulating layers to confine electrode and contact areas ofthe device. In some implementations, for example, the layer includes ahydrogel layer. For example, the layer (e.g., the hydrogel layer) canalso be disposed over the at least some of the electrochemical sensingelectrodes (e.g., the working electrode) of the anodic and cathodicelectrode assemblies to assist the flow of the extracted ISF to theelectrochemical sensor. In some implementations, for example, theflexible substrate includes paper, fabric, silicone, or polyurethane.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A shows a schematic illustration of an exemplary wearableelectrochemical sensor platform for non-invasive analyte sensing.

FIG. 1B shows a photograph of an exemplary iontophoretic-sensing tattoodevice applied to a human subject.

FIG. 1C shows a schematic illustration of the timeframe of an exemplaryon-body implementation and the different processes involved in eachphase.

FIG. 1D shows an illustration of the sweat inducing mechanism using thepresent technology.

FIG. 2A shows a data plot showing exemplary chronoamperometric responsedata of an exemplary wearable electrochemical sensor to increasingglucose concentrations, from 0 μM (dash) to 100 μM (plot ‘l’) in bufferin 10 μM increments.

FIG. 2B shows a data plot depicting exemplary interference data in thepresence of 50 μM glucose (plot ‘a’), followed by subsequent 10 μMadditions of ascorbic acid (plot ‘b’), uric acid (plot ‘c’) andacetaminophen (plot ‘d’).

FIG. 3 shows exemplary amperogram data plots obtained for non-invasiveglucose detection obtained from two human subjects, wearingsimultaneously an exemplary wearable electrochemical sensor for glucosedetection (A, A′) with and (B, B′) without the IP operation; andexemplary data plots (C,C′) showing the correlation between dataobtained from exemplary biosensors, with and without the IP procedure,and that obtained using a blood glucose (BG) meter.

FIG. 4 shows a data plot showing the combined exemplary data obtainedfrom exemplary wearable glucose sensors (plots ‘a_(i)’, ‘a_(ii)’ and‘a_(iii)’), No-GOx sensors (plots ‘b_(i)’, ‘b_(ii)’ and ‘b_(iii)’) andNo-IP sensors (plots ‘c_(i)’, ‘c_(ii)’ and ‘c_(iii)’) before and aftermeal consumption.

FIG. 5 shows a data plot depicting exemplary blood glucose levelsmeasured for two subjects, in which the subjects finished consuming foodat time T=0 min.

FIG. 6A shows a diagram of an exemplary non-invasive glucose sensingdevice of the disclosed technology including a non-invasive epidermalelectrochemical sensor and integrated processing and communicationsunit.

FIG. 6B shows a block diagram of an exemplary integrated processing andcommunications unit.

FIG. 7 is a schematic of an exemplary tattoo- or patch-basedelectrochemical biosensor device with different components of magneticcontacts.

FIGS. 8A and 8B are photos an exemplary glucose sensor patch device andexemplary flexible PCB electronics.

FIG. 9 shows exemplary results for three different subjects obtained bythe disclosed glucose sensor patch device in conjugation with thewireless electronic board.

FIG. 10 shows exemplary data from a separate study where the disclosedglucose sensor patch device recorded the glucose spikes occurring aftera subject consumed breakfast, lunch and snacks.

FIG. 11 shows an exemplary tattoo-based transdermal alcohol sensordevices and systems.

FIG. 12 shows exemplary results obtained using the disclosednon-invasive biosensor device implemented as a tattoo- or patch-basedalcohol sensor.

FIGS. 13A-B represent amperograms obtained using the disclosednoninvasive alcohol sensing from human subjects wearing the disclosedalcohol tattoo sensor.

FIG. 13C shows exemplary Amperograms obtained for noninvasive alcoholdetection obtained from human subjects wearing the alcohol tattoo sensorshowing correlation between BAC level and current response from tattoobiosensor measured before (i, BAC: 0%) and after drinking 5 oz. of wine(ii, BAC: 0.0025%)) and 10 oz. of wine (iii, BAC: 0.0062%)).

FIG. 13D shows resulting linear plot between current response and BAClevel.

FIG. 13E shows integration of flexible wireless circuit board withalcohol tattoo sensor.

FIG. 14 shows an exemplary Pilocarpine Induced Sweat Generation using anexemplary alcohol tattoo sensor device.

FIG. 15 shows an exemplary on-tattoo electrode.

FIG. 16 shows exemplary data on alumina SPE_EtOH detection.

FIG. 17 shows exemplary data on EtOH detection on Alumina substrate (PGdesign).

FIG. 18 shows exemplary data on EtOH detection on Tattoo paper.

FIG. 19 shows data for exemplary EtOH detection with Std.3 design ontattoo (PG stencil).

FIG. 20 shows data for exemplary EtOH detection with Std.3 design ontattoo (PG stencil).

FIG. 21 shows an exemplary tattoo substrate.

FIG. 22 shows exemplary electrode design and system for 2E and 3Esystems.

FIG. 23 shows data from an exemplary glucose tattoo electrode with 3Esystem (Silver counter and ref electrode.)

FIG. 24 shows data obtained from an exemplary PG tattoo Electrode with3E system (Prussian Blue carbon electrode as the counter electrode).

FIG. 25 shows schematic diagrams and real images displaying an exemplarybiodevice and enzyme and electrode reactions involved in the response ofgraphite-Teflon-AOD-HRP-ferrocene.

FIG. 26 shows an exemplary amperometric trace measured with an exemplarybiodevice at 0.00V depicted for a 5.0×10⁻⁴ molL⁻¹ ethanol solution.

FIG. 27 show exemplary current-time recording obtained after alcoholingestion by placing a biodevice on the skin: continuous mode recordingfor three different volunteers and single measurements carried out forone volunteer.

FIG. 28 shows exemplary correlation data between the cathodic currentvalues measured with an exemplary biodevice.

FIG. 29 shows exemplary least square straight line regression data.

FIG. 30 shows a comparison data of the BAC values obtained with thebiodevice through sweat measurements and with the gas chromatographymethod.

FIG. 31A shows a process diagram illustrating an exemplary fabricationmethod to produce epidermal electrochemical sensors of the disclosedtechnology, e.g., such as the exemplary T3 electrochemical sensors.

FIG. 31B shows images of exemplary T3 electrochemical sensors showingseveral exemplary printed designs.

FIG. 31C shows an image of an exemplary three-electrode T3 biosensorapplied to porcine skin for physiological monitoring.

FIG. 31D shows a schematic illustration of exemplary material layers ofan exemplary epidermal electrochemical sensor device.

FIG. 31E shows a block diagram of an exemplary embodiment of anepidermal electrochemical sensor device capable of being worn on skin ora wearable item.

FIG. 32A shows images for transferring an exemplary epidermalelectrochemical sensor device, e.g., such as the exemplary device shownin the images, on a user's skin.

FIG. 32B shows images of several representative design permutationstransferred onto the epidermis.

FIG. 32C shows images that validate the structural resiliency of theexemplary T3 sensors to extreme mechanical deformations, e.g., in whichvarious strain permutations were applied to the sensors.

FIG. 33A shows cyclic voltammogram data plots that were obtained for 2.5mM ascorbic acid (AA) at an exemplary SPE, at an exemplary T3 sensor onGORE-TEX, and at an exemplary CF-reinforced T3 sensor on GORE-TEX.

FIG. 33B shows cyclic voltammetric response plots that were obtained forthe detection of 2.5 mM uric acid (UA) at an exemplary SPE an exemplaryCF-reinforced T3 sensor on porcine skin.

FIG. 33C shows SWV plots that were obtained for the electrochemicaldetection of 225 μg/mL TNT at an exemplary SPE, an exemplaryCF-reinforced T3 sensor on porcine skin.

FIG. 34A shows a cyclic voltammogram plot illustrating the enhancedresponse generated by the dispersion of CF segments into the ink matrix,e.g., in which the scan rate was 10 mV/s.

FIG. 34B shows a resistive profile plot of a normal (black squares) andcarbon fiber-reinforced (red dots) 1 cm Ag/AgCl tattoo trace on porcineskin.

FIG. 34C shows a plot of Nyquist complex-valued impedance curvesgenerated by an exemplary T3 sensor before bending (black squares) andafter 10 bending operations (red dots) on porcine skin, e.g., in whichpotassium ferricyanide (K₃Fe(CN)₆) was employed as the redox probe.

FIG. 35A shows an IV data plot showing the effect of repetitive washingcycles upon the CV waveform generated at the tattoo biosensor (onporcine skin) using 2.5 mM UA.

FIG. 35B shows an IV data plot displaying the response of the exemplaryCF-reinforced T3 sensor to repetitive pinching operations employing 2.5mM AA as a redox probe.

FIG. 35C shows images captured immediately following the application ofthe T3 sensors on skin for an exemplary T3 sensor and an exemplaryCF-reinforced T3 sensor, respectively, and images captured after 18hours of continuous epidermal wear of the exemplary T3 sensor and theexemplary CF-reinforced T3 sensor, respectively.

FIG. 36A shows a process diagram illustrating a fabrication method toproduce epidermal electrochemical sensors with ion-selective electrodes.

FIG. 36B shows an image of an exemplary ISE tattoo sensor including twoelectrodes, e.g., including an ISE and a reference electrode, andconnection points that can interface with a voltmeter, for example, viaelectrically conductive conduits.

FIG. 37A shows a data plot of the potential-time response of anexemplary ISE tattoo sensor for decreasing pH levels and an inset plotof electrical potential versus pH, e.g., using the standard McIlvaine'sbuffers.

FIG. 37B shows a data plot of the potential-time response of theexemplary ISE tattoo sensors, which demonstrates the reproducibility ofthe sensors in response to large pH fluctuations.

FIGS. 38A and 38B show a data plot and images representing the influenceof repeated mechanical strain (e.g., bending) upon the response of anexemplary tattoo ISE device.

FIGS. 39A and 39B show a data plot and images representing the influenceof repeated mechanical strain (e.g., stretching) upon the response of anexemplary tattoo ISE device.

FIG. 40A shows a data plot of the real-time voltage-time response of anexemplary ISE tattoo sensor applied to a subject's neck to detect pHchanges, e.g., as compared to that of a conventional pH meter.

FIG. 40B shows a data plot of the real-time voltage-time response of anexemplary ISE tattoo sensor applied to a subject's wrist to detect pHchanges, e.g., as compared to that of a conventional pH meter.

FIG. 40C shows a data plot of the real-time voltage-time response of anexemplary ISE tattoo sensor applied to a subject's lower back to detectpH changes, e.g., as compared to that of a conventional pH meter.

FIG. 40D shows an image showing the exemplary device used in theseexemplary implementations (e.g., the tattoo ISE sensor interfaced with adigital multimeter) attached to the subject's wrist for the epidermalmeasurements of pH in human perspiration.

FIG. 41A shows a schematic illustration of an exemplary enzymatic T3electrochemical sensor device including three electrodes (e.g., theworking electrode, counter electrode, and reference electrode)configured in the design ‘tattoo’ design “NE” for electrochemicaldetection of L-Lactic acid.

FIG. 41B shows a schematic illustrating an exemplary modified workingelectrode including the transducer layer coated by biocompatible polymer(e.g., chitosan).

FIG. 42A shows an amperometric data plot of the responses for differentconcentrations of L-Lactate using the exemplary lactate T3electrochemical sensor device, e.g., with 1 mM increment,E_(applied)=+0.05V.

FIG. 42B shows the exemplary corresponding calibration plot ofL-Lactate.

FIG. 43A shows a data plot showing the stability of an exemplary lactateT3 electrochemical sensor.

FIG. 43B shows a data plot of an interference implementation with (a) 4mM L-lactate, (b) 84 μM creatinine, (c) 10 μM ascorbic acid, (d) 0.17 mMglucose and (e) 59 μM uric acid.

FIGS. 44A-44D show data plots of the electrochemical responses of anexemplary lactate T3 sensor transferred on a flexible GORE-TEX textileundergoing repeated bending (FIG. 44A) and stretching (FIG. 44C), withtheir normalized current plots (FIGS. 44B and 44D, respectively).

FIG. 45 shows images of exemplary lactate T3 sensors on human skin onthe neck under mechanical strain including stretching (top row), bending(middle row), and twisting (bottom row) endured by a bare ‘NE’ tattooduring and subsequent to 100 stretching, bending, and twistingiterations (shown in the right column of images).

FIG. 46 shows a schematic illustration of an exemplary method tofabricate tattoo biofuel cells using screen printing techniques.

FIG. 47 shows images of the epidermal tattoo biofuel cells duringmechanical stress caused by continuous body movements including (i)stretching, (ii) bending, and (iii) twisting.

FIG. 48A shows an image of the exemplary functionalized device.

FIG. 48B shows a data plot of polarization curves of the exemplaryfunctionalized MWNTs/TTF/LOx bioanode in the absence of presence of 14mM lactic acid in 0.2 M McIlvaine buffer solution, pH 5.5, respectively.

FIG. 48C shows a data plot of power density achieved from the exemplarytattoo biofuel cell device with different lactic acid concentrations.

DETAILED DESCRIPTION

Several diseases and disorders require patients to continuously monitorcertain analytes associated with the disease on a continuous basis,e.g., such as multiple times per day. Diabetes is one of the most widelyspread modern lifestyle diseases affecting hundreds of millions ofpeople and is among the leading causes of deaths globally. Patients withdiabetes must frequent monitor their glucose levels in order to managethe disease and avoid its associated problems, which can be severe anddeadly.

Extensive research and development of analyte monitoring devices hasevolved over the past few decades, which has led to the introduction andwidespread use of self-testing blood glucose meters. However, suchself-testing methods rely on inconvenient and painful blood samplingfrom the finger tip that compromises the patient's compliance.Furthermore, it requires active patient participation and this ‘patientinvolvement’ is a major factor affecting optimal disease management.

The present self-testing and self-monitoring blood glucose meter marketrelies on painful finger stick processes to draw blood samples followedby amperometric detection of glucose. While these finger stick-blooddrawing technologies can be highly accurate in detecting glycemic levelsin blood, the painful finger pricking is a major drawback. The need foractive patient participation can lead to compromised diabetesmanagement. For example, such finger stick (needle) devices have beenattributed to patient noncompliance, particularly among neonatals,children, elderlies, hemophobics and trypanophobics patients. Because itcan be hard for users to take blood samples, it is of great interest todevelop new ways to obtain blood samples without a skin puncture.

Some enzyme-based micro-needle sensors inserted under the skin tomeasure chemical analytes in tissue fluid have been introduced in themarkets to address these limitations. These micro-needle sensors are‘minimally’ invasive, as they include needles that pierce the human skinto detect glucose levels in the interstitial fluid (ISF). Theseminimally invasive sensing methods are based on the correlation betweenchemical concentration in the ISF and in blood. Yet, these minimallyinvasive devices too face limitations, for example, biofouling, fingerstick validation, and microbial infection. There have been severalprototypes of non-invasive glucose monitoring systems to overcome suchchallenges. However, low specificity towards glucose and skin irritationhas resulted in the withdrawal of some of these non-invasive devicesfrom the market.

Techniques, systems, and devices are described for non-invasive wearableelectrochemical sensors for detection of physiologically relevantchemicals for health and performance monitoring. The monitoring canoccur in continuous or quasi-continuous mode through multiplemeasurements performed using the same sensor over a time period (longerthan measured in minutes, e.g., a few hours to several days). Acontinuous chemical monitoring device can be obtained by selecting ahydrogel layer that allows fast transport of all the extracted chemicalanalyte to the sensor surface. Ideal hydrogels can be prepared by usingpolymers belong to the family of agarose, polyvinyl alcohol, acrylates,polyurethane, polyethylene glycol, etc. A combination of the abovepolymers can also be utilized to achieve the desired hydrogels. Thehydrogels can be stabilized (to avoid drying) by incorporatinghumectants, such as, glycerol, glycols, sugar polyols. By playing withthe ratio of the above components, the porosity of the hydrogel can bevaried, thus manipulating the transport speed of the analyte.

In some implementations, a control working electrode can be incorporatedinto the device to reduce the effect of interference and obtain highlyreliable data. The additional working electrode is not functionalizedwith the receptor (e.g.: enzyme) but has all the other components. Thisadditional electrode can be placed close to the receptor functionalizedworking electrode. Data can be recorded for both these electrodessimultaneously and compared to obtain interference-free signal (Signalfrom control working electrode is subtracted fromreceptor-functionalized electrode).

The transport speed in the ideal hydrogel can depend on a number offactors including hydrogel porosity, iontophoresis current, etc. Anideal hydrogel remains hydrated for at least a week under ambientconditions and allows rapid mass transport, thus reducing the timerequired for iontophoresis and detection to a few minutes (<5 min).Furthermore, the high mass transport speed towards the working electrodeallows rapid detection and consumption of most of the extracted analyteto obtain a analyte-free hydrogel layer, ready for the nextiontophoresis-detection cycle without any memory effect (sensor responsebeing affected by the analyte extracted during previousiontophoretic-detection cycle). A rough estimate for fast analytetransport that allows the analyte to move quickly from the skin to theelectrode surface can be approximately 0.1 mm/s. The higher the speed,the faster the transport and the better the result.

In some implementations, the hydrogel can be optimized to limitbiofouling and enhance sensor stability. For example, the hydrogel canalso incorporate anti-fouling, bactericidal and anti-fungal chemicalagents widely used in the wound dressing industry.

For example, the disclosed technology can be implemented as devices,systems and methods for a wearable and non-invasive glucose sensor forcontinuous monitoring of glucose levels for efficient diabetesmanagement. In some aspects, the wearable glucose sensor device includesa reverse iontophoretic-amperometric hybrid platform on diverse wearablesubstrates to obtain rapid and sensitive glucose response, in whichglucose is extracted from the ISF to the skin surface by passing a lowcurrent through the human skin followed by detection of the extractedglucose using a highly selective amperometric sensor. For example, insome embodiments, the wearable substrates include temporary tattoo paperand stretchable elastomeric membranes that attach to the skin forcontinuous glucose monitoring without the need for blood sampling.

The present electrochemical sensor technology includes single-use‘place-detect-dispose’ (PDD) on-body sensors for non-invasive andpainless detection of particular analytes associated with disease orhealth performance metrics. In some aspects, the single-use mode of thedisclosed wearable electrochemical sensors can be employed on a devicepatch that can be fabricated or integrated on various substrates, e.g.,including, but not limited to, paper, fabric, bendable and/orstretchable plastics or stretchable elastomeric membranes. Exemplarysingle-use wearable electrochemical sensor devices of the presenttechnology can be fabricated using variety of techniques, e.g.,including, but not limited to, printing processes such as screenprinting, roll-to-roll printing, ink-jet printing, and/or lithographictechniques. The present technology can provide a pain-free single-usenon-invasive test as an alternative to single-use strips used fordetecting chemicals in blood.

Some implementations of the wearable electrochemical sensor devices canbe used for detecting chemicals in the ISF. In such applications, forexample, an electrochemical sensor device includes a reverseiontophoretic electrode assembly that will first extract the ISF to thesurface of the skin followed by electrochemical detection of the desiredchemical analyte via electrochemical sensor electrodes of the device.Other implementations of the wearable electrochemical sensor devices canbe used for detecting chemicals in a user's biological fluid, such assweat. In such applications, for example, the electrochemical sensordevice includes an iontophoretic electrode assembly that will firstinduce localized sweating (in the vicinity of the electrochemicalelectrodes platform the device) by transdermally administeringsweat-inducing drugs into the skin, followed by detection of thechemical analyte in the induced sweat via the electrochemical sensorelectrodes.

The disclosed wearable electrochemical sensor devices and techniques canbe used for diabetes management by providing a non-invasive highlyselective single-use pain-free glucose monitoring platform, e.g., as analternative to the painful strip-based glucose monitors. The disclosedtechnology can also be used for detecting and analyzing other bloodanalytes, for example, such as for non-invasive detection of alcohol(e.g., in which a test can be utilized by the user in car or bar beforedriving), lactate (e.g., in sport activity) or urea (e.g., for kidneypatients), among others.

In one exemplary embodiment, a non-invasive electrochemical sensordevice of the disclosed technology includes an anodic and a cathodicelectrode contingent. Each contingent includes an electrochemical sensorelectrode assembly including working electrode, a counter and/orreference electrode, and an iontophoresis electrode. The iontophoreticelectrodes are configured to not be part of the electrochemical sensorelectrode assembly. The iontophoretic electrodes operate independentlyfrom the electrochemical sensor electrode assembly. The iontophoreticelectrodes are utilized for extraction of ISF to the surface of the skinor for administrating sweat-inducing drug for generating localizedperspiration. The working, and counter and/or reference electrodestogether constitute the electrochemical sensor electrode assembly thatis utilized for sensitive chemical detection. In implementations of atwo electrode configuration of the electrochemical sensor, for example,in addition to the working electrode, the second electrode is operableas a reference and counter electrode. In implementations of a threeelectrode configuration of the electrochemical sensor, for example, inaddition to the working electrode, one electrode is operable as areference electrode, and the other electrode is operable as a counterelectrode. In some embodiments, for example, the working electrode ofthe cathode and/or the anode is modified with specific receptors likeenzyme, ionophores and/or other reagents for achieving selectivedetection of the desired chemical analyte. In some implementations, forexample, the anodic and cathodic contingent can be coated with abiocompatible hydrogel layer, e.g., covering the iontophoreticelectrodes, or covering the iontophoretic electrode and at least some ofthe electrochemical sensor electrodes. For example, the biocompatibleand conductive hydrogel layer can be disposed over the electrodes of theanodic and cathodic contingent to assist the flow of the extracted ISFto the electrochemical sensor. In some implementations, for example, thehydrogel layer can be loaded with specific chemicals agents (e.g.,including FDA approved chemical agents) to enhance ISF extraction at lowiontophoretic currents (in case of chemical detection in ISF) or withsweat-inducing drugs for localized sweating (in case of chemicaldetection in sweat).

The non-invasive electrochemical sensor device, when attached to theskin and electrically coupled to one or more electrical circuits orelectronic devices, is operable to detect the chemical analyte from theextracted ISF or the sweat induced to the outer or local environment ofthe skin. For example, the non-invasive electrochemical sensor devicecan include an electrode interface assembly comprising individual orindependent electrically conductive conduits disposed on the flexiblesubstrate that are electrically coupled to the anodic and cathodicelectrode assemblies. The flexible substrate can include bendable and/orstretchable properties. The individual or independent conduits areconfigured to electrically couple the electrochemical sensor electrodesof the anodic and cathodic electrode assemblies and the iontophoreticelectrodes to the external electrical circuits or electronic devices,which are able to electrically energize the electrochemical sensingelectrodes and to separately electrically energize the iontophoreticelectrode for their respective operations.

In some implementations for alcohol detection, a three electrode systemfor the electrochemical sensor electrode assembly is preferred. Whereasfor glucose detection, the electrochemical sensor electrode assembly mayinclude a two electrode system (e.g., working electrode andcounter/reference electrode) because the current measured as a functionof glucose concentration is relatively low, such that the two electrodesystem is sufficient to detect glucose. For example, in case of alcoholdetection, the concentration of alcohol can be relatively high in thebiofluid (e.g., sweat), and hence the current measured is higher, andthereby a three electrode system can be utilized.

FIG. 1A shows a schematic illustration of an exemplaryiontophoretic-electrochemical sensor platform 100 for non-invasiveanalyte sensing, e.g. such as for glucose as shown in FIG. 1C, includinga printable iontophoretic-sensing system displaying the tattoo-basedpaper (e.g., substrate shown in purple 102), Ag/AgCl electrodes (e.g.,electrode contingents shown in silver 104), Prussian Blue electrodes(e.g., transducer shown in black 106), transparent insulating layer(e.g., shown in green 108) and hydrogel layer (e.g., shown in blue 109).FIG. 1B shows a photograph of an exemplary glucoseiontophoretic-biosensing device applied to a human subject.

To operate the device, a wearer applies the non-invasive electrochemicalsensor device to his/her skin for non-invasive chemical detection.

During operation for detection of the chemical analyte or analytes inISF, a mild current (e.g., such as 0.3 mA/cm² or less) is passed betweenthe iontophoretic electrodes (e.g., Ag/AgCl electrodes) of the anodicand cathodic electrode contingents to create an electric field throughthe wearer's skin to extract ISF to the surface of the skin by drivingions of the ISF toward the electrode contingents. For example, theelectric field can be applied for one, a few, or several minutes (e.g.,5 to 10 minutes) to extract sufficient ISF containing physiologicallyrelevant and detectable glucose levels to the surface of the skin.Subsequently, the chemical analyte present in the extracted ISF isquantified by the electrochemical sensor. For example, theelectrochemical sensor electrodes can be operated to performamperometric, potentiometric or voltammetric techniques that produce anelectrical signal at the sensor electrodes, e.g., associated with anelectrochemical or redox reaction sustained at the anodic and cathodicelectrode contingents. Operation of the non-invasive electrochemicalsensor device preferably includes extracting the ISF prior to operatingthe electrochemical sensor electrodes for electrochemical analysis,e.g., to save power and to avoid interference in the analyte detectionthat could be caused by the iontophoretic electrodes. However, in someimplementations, the extracting and the detecting techniques can beperformed concurrently.

During operation for detection of the chemical analyte or analytes in auser's sweat, a mild current (e.g., such as 0.3 mA/cm² or less) ispassed between the iontophoretic electrodes to create an electric fieldto administer sweat-inducing drug or chemical compounds into the skin.For example, the drugs or chemical compounds can be stored in theoverlaying hydrogel over one or both of the iontophoretic electrodes(e.g., such as the hydrogel over the iontophoretic electrode of theanodic contingent). When the iontophoretic current passes through thehydrogel layer and through the skin, the electric field creates a forceupon the drugs or chemical compounds that drives them into the skin. Thedrug or compound, while in the skin, induces the physiological processof perspiring (sweat) in the location of the drug or compoundadministration. FIG. 1D shows an illustration of the sweat inducingmechanism using the present technology. For example, the drug may be apositively charged compound so that electrostatic repulsion between theanode and the drug causes the drug to be administered into the skin. Inthe example shown in FIG. 1D, pilocarpine (Pilo) is illustrated.Subsequently, the chemical analyte present in the sweat that is presentdue to the sweat-inducing mechanism is quantified by the electrochemicalsensor. For example, the electrochemical sensor electrodes can beoperated to perform amperometric, potentiometric or voltammetrictechniques that produce an electrical signal at the sensor electrodes,e.g., associated with an electrochemical or redox reaction sustained atthe anodic and cathodic electrode contingents.

Upon completion of the test, the user can remove or replace the deviceand subsequently dispose of the removed device. For the next test, awearer may utilize a new non-invasive electrochemical sensor device.

The non-invasive electrochemical sensor device is connected to anexternal electronic system, which can be portable and/or wearable on theuser. The electronic system is used to power the electrochemical sensordevice and analyze the acquired sensor signal (e.g., detected electricalcurrent, voltage, etc.) to produce data on the analyte, e.g., such asthe chemical concentration. In some implementations, the electronicsystem can include a display to present the analyzed data to the user.The information can be transmitted wirelessly from the electronic systemto a user computing device, e.g., such as a smartphone, tablet, wearablecomputing device such as smartglasses, smartwatch, etc., and/or laptopor desktop computer. The electronic system can be connected to such usercomputing devices via physical contact (e.g., wired) or wirelessly usingRF or Bluetooth communication, or other wireless communicationtechniques.

FIGS. 1A and 1B show a diagram and image of an exemplary non-invasiveelectrochemical sensor device for skin-worn glucose sensing on atemporary ‘tattoo’ platform, which can be operated by the described‘place-detect-dispose’ on-body monitoring technique. The example deviceshown in FIGS. 1A and 1B can be utilized for detecting other analytes inISF (e.g., lactate, urea, etc.) and sweat (e.g., alcohol, electrolytes,etc.). Also, such devices can be fabricated or integrated on otherplatforms, for example, paper, fabric, flexible plastics or stretchableelastomeric membranes using variety of fabrication routes. For example,the substrate can include silicone or polyurethane membranes, e.g., suchas a polyurethane-based elastomer. The example device shown in FIGS. 1Aand 1B include substrates with highly favorable substrate-skinelasticity with an attractive electrochemical performance.

FIG. 1C shows an illustrative diagram of using the exemplarynon-invasive electrochemical sensor device for glucose monitoring. Theexample device uses a lower current density to extract the ISF glucosefollowed by selective amperometric biosensing using a glucose oxidase(GOx)-modified Prussian Blue transducer. Such flexible, low-cost andaesthetically pleasing iontophoretic-biosensing platform can be easilymated with the human skin with minimal intrusion to the wearer'sroutine. For the extraction and sensing operations using such printableskin-worn platform, the example device includes a Ag/AgClreverse-iontophoresis electrodes (which can include an agarose hydrogelcoating) for the anodic and cathodic electrode contingents, which areoperable to efficient deliver ISF close to the working andcounter/reference electrodes of the electrochemical sensor assemblies ofthe anodic and cathodic electrode contingents, as shown in FIGS. 1A and1C. The working electrodes can include a biocatalytic reagent layer thatcan be optimized for imparting the sensitivity needed for detecting low(micromolar) glucose concentrations in the extracted ISF and highspecificity in the presence of common interfering electroactive species.

An example of the non-invasive electrochemical sensor device for glucosemonitoring includes a non-invasive wearable continuous glucosemonitoring device for pain-free diabetes management. A continuouschemical monitoring device can be obtained by selecting a hydrogel layerthat allows fast transport of all the extracted chemical analyte to thesensor surface. The device can be fabricated on user-friendlysubstrates, e.g., like temporary tattoo paper/stretchable elastomericmembrane, and thus causes least levels of intrusion to the wearer. Thedevice relies on extraction of glucose from the ISF to the skin surfaceby passing a low current through the human skin followed by detection ofthe extracted glucose using a highly selective amperometric sensor. Thedevice uses low current, which at least minimizes skin irritation. Insome implementations, for example, the current density can be furtherlowered by incorporating skin permeation enhancers within the system.

An example of the protocol for continuous non-invasive glucose sensorincludes applying of the device to the target skin location, such as,deltoid followed by turning ON the device. Upon activating the device,the reverse iontophoretic process is activated for a few minutes (<5min) to extract ISF followed by deactivating the reverse iontophoreticprocess and activating the detection process to detect the extractedchemical analyte, for example, glucose. The iontophoretic and detectionprocesses are repeated every 20-30 minutes using the worn device duringa fixed period of time (24 hours to 1 week) to continuously monitorglucose levels. A similar process is also employed to detect otherchemicals.

In one exemplary embodiment, a skin-worn temporary-tattoo basednon-invasive glucose monitoring platform coupling an amperometricbiosensor with a reverse iontophoresis operation as shown in FIG. 1A.The exemplary glucose monitoring platform includes body-compliantwearable electrochemical devices based on temporary tattoos that combinehighly favorable substrate-skin elasticity with an attractiveelectrochemical performance. In some implementations, the skin-worntattoo-based glucose detection device.

During operation, a mild current is passed between the Ag/AgCliontophoretic electrodes through the wearer's skin to extract glucosemolecules present in the ISF to the surface of the skin, as depicted inthe illustration of FIG. 1C. The time and magnitude of the currentpassed can be further lowered by incorporating skin permeation enhancerswithin the hydrogel. Subsequently the extracted glucose is detectedusing a sensitive amperometric glucose sensor.

The wearable non-invasive sensor for detecting glucose has been utilizedin example implementations that included applying it to human subjectsto detect glucose spikes occurring due to food consumption. Suchexemplary testing has been carried out at a fixed time after the mealconsumption using the sensor in single-use ‘place-detect-dispose’ mode.The response obtained with said single-use glucose sensor has also beencorrelated with that obtained from a commercial blood glucose meter.

Exemplary implementations of the disclosed non-invasiveiontophoretic-electrochemical sensing device were performed for in vitrooptimization of the sensor sensitivity and selectivity, and innon-invasive glucose monitoring in human subjects, which was validatedby simultaneous blood finger stick measurements using a commercialglucose meter. For example, the specificity of the exemplary wearableiontophoretic-electrochemical GOx sensor was validated by applying itsimultaneously with an enzyme-free sensor (no GOx control) on humansubjects. In some implementations, other enzymes, such as glucosedehydronates (GDH) can be used in the non-invasive biosensor disclosedin this patent document. Reverse iontophoresis prior to detection wasdemonstrated by analyzing the sensor response with and without activeextraction of glucose ISF towards the sensor surface. The exemplaryon-body implementations reveal that the exemplaryiontophoretic-biosensing platform is capable of non-invasive glucosemonitoring in real-life situations. For example, attractive features ofthe disclosed skin-worn sensor technology also highlight its potentialfor on-body monitoring of other target chemicals present in theinterstitial fluid.

Exemplary reagents and instrumentation used in the exemplaryimplementations included the following. Glucose oxidase (GOx) fromAspergillus niger, Type X-S (EC 1.1.3.4), chitosan, bovine serum albumin(BSA), sodium phosphate monobasic (NaH2PO4), sodium phosphate dibasic(Na2HPO4), D(+)-glucose, L(+)-ascorbic acid, uric acid, acetaminophenand agarose were obtained from Sigma-Aldrich (St. Louis, Mo.). Aceticacid was obtained from EMD Chemicals Inc. (Gibbstown, N.J.). Allreagents were used without further purification. Electrochemicalcharacterizations were performed at room temperature using a CHInstruments electrochemical analyzer (model 1232A, Austin, Tex.) andPGSTAT 101 from Metrohm Autolab (Netherlands).

An exemplary fabrication process of the exemplary glucose sensorplatform was implemented, followed by modification and transferprocesses. For the exemplary implementations, sensor patterns weredesigned in AutoCAD (Autodesk, San Rafael, Calif.) and outsourced forfabrication on stainless steel through-hole 12 in.×12 in. framedstencils (Metal Etch Services, San Marcos, Calif.). Papilio temporarytransfer tattoo base paper was purchased from HPS LLC (Rhome, Tex.). Asequence of the Prussian blue conductive carbon (C2070424P2, GwentGroup, Pontypool, UK), silver/silver chloride (Ag/AgCl) ink (4001,Engineered Conductive Materials, LLC, Delaware, Ohio), and insulator(Dupont 5036, Wilmington, Del.) inks were patterned on the substrateemploying an MPM-SPM semiautomatic screen printer (SpeedlineTechnologies, Franklin, Mass.). As illustrated in FIG. 1A, theiontophoretic-electrochemical sensor design includes a pair of reverseiontophoresis electrodes (Ag/AgCl ink), a pseudo reference/counter(Ag/AgCl ink) and working electrodes (Prussian Blue ink). A transparentinsulator was screen printed over the surface of the electrode patternto confine the electrode and contact areas. The Ag/AgCl ink was cured at130° C. for 3 min, while the Prussian Blue ink was cured at 80° C. for10 min in a convection oven.

Following the printing of the exemplary electrode transducers, theworking electrode was functionalized with the reagent layer. The enzymeGOx solution (e.g., 34 mg/mL containing 10 mg/mL BSA stabilizer) wasmixed with chitosan solution (0.5 wt % in 0.1 M acetic acid) in 1:1 v/vratio. Subsequently, a 2 μL droplet of the above solution was casted onthe electrode and dried under ambient conditions.

Exemplary in vitro characterizations were performed using a 0.1 Mphosphate buffer (pH 7.0) solution containing 133 mM NaCl. The operatingpotential for the exemplary non-invasive wearable glucose sensor wasselected by using cyclic voltammetry. The amperometric response wasrecorded after 1 min incubation in the sample solution, using apotential step to −0.1 V (vs Ag/AgCl) for 60 s. The sensor specificitywas examined in the presence of relevant electroactive constituents,namely, 10 μM each of ascorbic acid, uric acid and acetaminophen.

Exemplary on-body glucose monitoring characterizations were performedusing the exemplary iontophoretic, non-invasive wearable glucose sensordevice. A continuous chemical monitoring device can be obtained byselecting a hydrogel layer that allows fast transport of all theextracted chemical analyte to the sensor surface. An agarose hydrogel,covering all the electrodes, was applied to the iontophoretic glucosesensor. For example, the hydrogel was prepared by heating a continuouslystirred agarose solution (4% w/v) in 0.1 M phosphate buffer (pH 7) at120° C. for 15 min. The solution was then cooled down to 60° C. and 100μL of the solution was casted on the sensor area to form a uniformhydrogel layer covering all the three electrodes of both the anodic andcathodic contingents. The epidermal biosensor evaluation was performedin strict compliance with a protocol approved by the institutionalreview board (IRB) at the University of California, San Diego. A totalof 7 consenting healthy volunteers (e.g., 4 males and 3 females betweenthe ages of 20 and 40), with no prior medical history of heartconditions, diabetes, or chronic skeletomuscular pain, were recruitedfor participation in the study. The subjects were requested to arrive atthe lab in fasting state. The epidermal studies comprised oftransferring the iontophoretic glucose sensor to the skin followed byapplying a constant current of 0.2 mA/cm² between the tworeverse-iontophoresis electrodes for 10 min to extract ISF to thesurface of the skin and finally recording the amperometric glucoseresponse at an applied potential of −0.1 V (vs Ag/AgCl) for 5 min. Thereverse-iontophoresis/detection cycle was performed first in the fastingstate followed by consumption of a carbohydrate-rich meal. Thereafter,each subject was requested to wait for 5 min before a similarreverse-iontophoresis/detection cycle was repeated to measure thepost-meal sensor response. The exemplary procedure is shownschematically in FIG. 1C. The crucial role of reverse iontophoresis wasexamined by analyzing the response obtained from two exemplary glucosebiosensors (applied simultaneously to subjects) with and without reverseiontophoresis. The selectivity of the on-body sensor to glucose wasevaluated using two exemplary sensors, one containing the GOx enzymewhile the other devoid of it, applied simultaneously on the subjects'deltoid. For each human trial, simultaneous finger stick blood glucosemeasurements were performed using commercial glucose strips (e.g.,Accu-Chek Aviva Plus®) to establish the correlation between the responseof the exemplary iontophoretic glucose sensor device and that obtainedfrom the commercial glucose meter.

The disclosed iontophoretic-biosensing device includes a differentelectrode pattern that includes the iontophoretic electrodes, e.g., ascompared to 3-electrode designs. The glucose sensor of the of theexemplary iontophoretic-biosensing device shown in FIG. 1A includes theanodic and cathodic contingents, in which each contingent includes: anAg/AgCl electrode that operates as a counter/reference electrode; aprintable Prussian-Blue transducer (e.g., which was selected in thisexemplary embodiment in view of its high selectivity towards hydrogenperoxide, the detectable product of the GOx enzymatic reaction); and anadditional Ag/AgCl reverse iontophoretic electrode which encompassed theworking and the counter/reference electrodes for efficient extraction ofISF close to the working and counter/reference electrodes. During thereverse iontophoresis operation, glucose is extracted at the cathodiccontingent, and the working electrode of the cathodic contingent (e.g.,modified with the GOx enzyme) can operate to selectively detect glucose.In other implementations where the analyte include a negative charge,e.g., such as lactate, the analyte is extracted at the anodiccontingent, and the working electrode of the anodic contingent can bemodified with the agent (e.g., catalyst, such as lactate oxidase (LOxenzyme) in the case of lactate) to selectively detect the analyte. Inthe exemplary implementations performed for glucose detection, chitosanwas utilized as a polymeric matrix for immobilizing the enzyme on thetransducer surface. However, in some implementations, a differentbiocompatible polymer other than chitosan can be used. While performingreverse iontophoresis, care should be taken to ensure proper contactbetween the skin and the sensor for efficient glucose extraction and toavoid skin irritation. For example, this can satisfied by evenly coatinga layer of biocompatible agarose gel on each contingent to cover all theelectrodes. The resulting glucose sensor can be easily applied to theskin, adhering and conforming to the contours of the epidermis, similarto a typical rub-on temporary tattoo, e.g., as shown in FIG. 1B.

Exemplary Results of In Vitro Implementations

The glucose level in the ISF is in the same concentration range as thatin the blood. However, the concentration of the ISF glucose extractedvia reverse iontophoresis is approximately two orders lower than that ofthe corresponding ISF glucose level. Keeping this in view the responseof the exemplary glucose tattoo sensor was evaluated over the 0-100 μMglucose concentration range, shown in FIG. 2A. These well-definedchronoamperometric responses to 10 μM glucose additions (a-l) revealedthat the sensors responded linearly and favorably in this range, andcould thus be utilized for detecting relevant ISF glucose levelsextracted during on-body applications. Specificity of a sensor is ofutmost importance for avoiding false alarms. Hence, the effect ofphysiologically relevant concentrations of common co-existinginterfering electroactive species on the sensor response was examined.The exemplary results, displayed in FIG. 2B, highlight the highspecificity of the sensor towards glucose (FIG. 2B, plot ‘a’) inpresence of ascorbic acid, uric acid and acetaminophen (FIG. 2B, plot‘b’-‘d’). Overall, the high sensitivity and selectivity demonstrated inFIGS. 2A and 2B reflect the coupling of the specific biocatalyticreaction with the low-potential amperometric transduction at theexemplary Prussian-Blue transducer, e.g., as compared to the highdetection potential utilized in GlucoWatch® that could lead tocompromised selectivity.

FIG. 2A shows a data plot showing exemplary chronoamperometric responsedata of an exemplary glucose sensor to increasing glucoseconcentrations, from 0 μM (dash) to 100 μM (plot ‘l’) in buffer in 10 μMincrements. FIG. 2B shows a data plot depicting exemplary interferencedata in the presence of 50 μM glucose (plot ‘a’), followed by subsequent10 μM additions of ascorbic acid (plot ‘b’), uric acid (plot ‘c’) andacetaminophen (plot ‘d’). In these plots, the exemplary conditionsincluded: potential step to −0.1 V (vs Ag/AgCl); medium ofphosphate-buffer with 133 mM NaCl (pH 7).

Exemplary Results of On-Body Implementations

After demonstrating in vitro the ability of the analyte sensors toselectively measure micromolar glucose levels, on-body detection of ISFglucose levels was examined in human subjects under real-life scenarioswith the system worn over the skin. Meal consumption triggers a rapidrise in blood glucose levels that may lead to detrimental effects ondiabetic patients. Hence, the present technology provides the ability ofthe non-invasive tattoo-based sensor to monitor such sudden glycemicspikes. In the exemplary implementations, the first task was to identifythe most appropriate time to perform the post-meal glucose sensing.Post-meal blood glucose levels of two subjects (1 male and 1 female)were thus measured at 10 min intervals over a 1 hour period followingcarbohydrate-rich meal (shown in FIG. 5). Based on these findings andliterature data indicating approximately 15-20 min lag time between ISFand blood glucose levels, a 5 min waiting period (followed by 10 min ofIP extraction) was considered for the post-meal glucose sensing on 7human subjects. None of the subjects reported perceptible discomfortduring these on-body studies. Only a mild tingling feeling at the skinunder the iontophoresis electrodes was experienced by few subjects forless than 10 s at the beginning of the test.

Two example control experiments were carried out to corroborate thevalidity of the reverse iontophoresis-based glucose sensing system: (1)detection of passively diffused ISF glucose by a GOx-modified sensor(No-IP sensor), and (2) use of an unmodified (enzyme-free) sensor underactive reverse-iontophoretic extraction of ISF (No-GOx sensor). Subjectswere selected randomly to participate in each set of controlexperiments. For each subject the control sensor was applied adjacent toa glucose sensor on the deltoid with a spatial gap of approximately 1.5cm. The response of the control sensor was recorded in tandem with theglucose sensor. FIG. 3 displays exemplary data obtained from twosubjects simultaneously adorning the glucose sensor and the No-IPsensor.

FIG. 3 shows exemplary amperogram data plots obtained for non-invasiveglucose detection obtained from two human subjects, wearingsimultaneously an exemplary glucose tattoo sensor (A, A′) with and (B,B′) without the IP operation; and exemplary data plots (C,C′) showingthe correlation between data obtained from tattoo biosensors, with andwithout the IP procedure, and that obtained using a blood glucose (BG)meter. Exemplary conditions included: potential step to −0.1 V (vsAg/AgCl).

As shown in FIG. 3, the data plot displays exemplary data obtained fromtwo exemplary subjects simultaneously adorning the glucose tattoo sensorand the No-IP sensor. It can be clearly noted that the respectiveglucose tattoo sensor display a distinct increment in the post-mealcurrent response (FIG. 3 A, A′ plot b) as compared to the fasting state(FIG. 3 A, A′ plots a). In contrast, for example, the respective No-IPsensors show minimal change in the current response before and after themeal (FIG. 3 B, B′ plots; a vs b). This exemplary study underpins theimportance of active reverse iontophoretic extraction of ISF glucose forperforming non-invasive glucose detection. Simultaneous blood glucosemeasurements using a commercial Accu-Chek Aviva Plus® glucose meter andcomparison with the response obtained from the tattoo sensors reveal thecorrelation between the non-invasive tattoo sensor and the blood glucosemeasurements.

Additional control experiments were carried out in the exemplaryimplementations for other subjects wearing the glucose tattoo sensoralong with a No-GOx sensor. In this set of exemplary studies, theresponse from glucose tattoo sensors was also significantly highercompared to that of the enzyme-free sensors, highlighting thespecificity of the sensor to detect the glucose substrate in presence ofpotential interfering species. FIG. 4 shows a data plot showing thecombined exemplary data obtained from glucose tattoo sensors (plots‘a_(i)’, ‘a_(ii)’ and ‘a_(iii)’), No-GOx sensors (plots ‘b_(i)’,‘b_(ii)’ and ‘b_(iii)’) and No-IP sensors (plots ‘c_(i)’, ‘c_(ii)’ and‘c_(iii)’) before and after meal consumption. Exemplary conditions fordata acquisition were those as in FIG. 3.

As shown in FIG. 4, the data plot displays exemplary a collection ofamperometric signals recorded with the glucose sensors (plots ‘a_(i)’,‘a_(ii)’ and ‘a_(iii)’), No-GOx sensors (plots ‘b_(i)’, ‘b_(ii)’ and‘b_(iii)’) and the No-IP sensors (plots ‘c_(i)’, ‘c_(ii)’ and ‘c_(iii)’)for different human subjects. These data clearly illustrate the abilityof the tattoo sensors to detect spikes in the glucose level occurringdue to food consumption. Another control experiment of the exemplaryimplementations was performed to identify the variation in the sensorresponse in absence of glucose spike. During this exemplary controlexperiment, a glucose tattoo sensor and a No-GOx sensor were appliedsimultaneously to a human subject. It was noted that both the bloodglucose level as well as the response from the two sensors remainedfairly stable, thus underscoring the sensor's ability to specificallydetect blood glucose spikes.

In some embodiments, the disclosed iontophoretic-biosensor devicesinclude an integrated electronic backbone for powering the sensor, andsignal processing and wireless communication units on the flexiblewearable sensor platform, capable of collecting continuous data from thediabetes patient, and performing large-scale glucose monitoring acrossdiverse patient populations. The exemplary iontophoretic-biosensingplatform can be readily used for the non-invasive monitoring of otherchemical markers present in the interstitial fluid, and fortranscutaneous drug delivery.

FIG. 6A shows a diagram of an exemplary non-invasive electrochemicalsensor of the disclosed technology including a non-invasive epidermaliontophoretic electrode assembly and electrochemical sensor electrodeassembly with an integrated processing and communications unit 600. FIG.6B shows a block diagram of an exemplary integrated processing andcommunications unit 600. As shown in FIG. 6B, the processing andcommunications unit 600 can include a power source 602 (e.g., abattery), a current injector 604, a potentiostat (sensor signalacquirer) 606, a data processing unit 608 capable of signal processingand communications (e.g., to external devices), and DC/DC and/oranalog-to-digital converters and/or signal conditioning circuits 610,612. The data processing unit can include a processor to process dataand a memory in communication with the processor to store data. Forexample, the processor can include a central processing unit (CPU) 614or a microcontroller unit (MCU). For example, the memory 616 can includeand store processor-executable code, which when executed by theprocessor, configures the data processing unit to perform variousoperations, e.g., such as receiving information, commands, and/or data,processing information and data, and transmitting or providinginformation/data to another entity or to a user. In someimplementations, the data processing unit can be implemented by acomputer system or communication network accessible via the Internet(referred to as ‘the cloud’) that includes one or more remotecomputational processing devices (e.g., servers in the cloud). Tosupport various functions of the data processing unit, the memory canstore information and data, such as instructions, software, values,images, and other data processed or referenced by the processor. Forexample, various types of Random Access Memory (RAM) devices, Read OnlyMemory (ROM) devices, Flash Memory devices, and other suitable storagemedia can be used to implement storage functions of the memory unit. Thedata processing unit can include an input/output unit (I/O) 618 that canbe connected to an external interface, source of data storage, ordisplay device. For example, various types of wired or wirelessinterfaces compatible with typical data communication standards can beused in communications of the data processing unit via the wirelesstransmitter/receiver unit 620, e.g., including, but not limited to,Universal Serial Bus (USB), IEEE 1394 (FireWire), Bluetooth, IEEE802.111, Wireless Local Area Network (WLAN), Wireless Personal AreaNetwork (WPAN), Wireless Wide Area Network (WWAN), WiMAX, IEEE 802.16(Worldwide Interoperability for Microwave Access (WiMAX)), 3G/4G/LTEcellular communication methods, and parallel interfaces. The I/O of thedata processing unit can also interface with other external interfaces,sources of data storage, and/or visual or audio display devices, etc. toretrieve and transfer data and information that can be processed by theprocessor, stored in the memory unit, or exhibited on an output unit ofan external device. For example, an external display device can beconfigured to be in data communication with the data processing unit,e.g., via the I/O, which can include a visual display device, an audiodisplay device, and/or sensory device, e.g., which can include asmartphone, tablet, and/or wearable technology device, among others.

In some embodiments, for example, the electronic system can be containedin a housing that electrically connects to the sensor device viaelectrical contact pads on the substrate of the sensor device that areinterconnected to the electrodes of the sensor device. In suchembodiments, the housed electronic system can be a portable device thatattaches and detaches from the device, and be stored on the user to bereadily available for the user's next test, e.g., such as in a user'spocket, purse, etc. The sensor device can make electrical contact withthe portable device via a number of connections including pressurecontacts, magnetic contacts, soldering contacts, etc. The housedelectronic system can be in wired or wireless connection with a user'smobile communication or computing device, e.g., such as a smartphone,tablet, wearable computing device such as smartglasses, smartwatch,etc., and/or laptop or desktop computer. The exemplary housed electronicsystem can supply power, operate, and retrieve the acquiredphysiological-related electrical signals from the sensor device.

Tattoo- or Patch-Based Wearable Electrochemical Biosensor Device

In some implementations, the disclosed technology can be implemented asa skin-worn tattoo- or patch based wearable electrochemical biosensordevice for non-invasive analyte monitoring, including glucose andalcohol. Exemplary implementations using an exemplary flexible tattoo-or patch-based glucose biosensor device of the present technology caninclude in vitro characterization of the tattoo sensors that showedtheir ability to detect micromolar levels of glucose in the presence ofcommon interfering chemical species; and including on-body evaluation ofthe tattoo-based iontophoretic-biosensing platform that showed theability to detect the rise in the glucose level after a meal in anon-invasive fashion. The disclosed tattoo iontophoretic-biosensingplatform can be utilized for everyday use in diabetes management. Thepresent technology demonstrates a non-invasive highly selectivepain-free continuous glucose monitoring platform. The exemplarywear-and-forget device can be easily worn by a person to autonomouslymonitor glycemic levels, e.g., thereby avoiding the need for activepatient involvement. The wearable device also has immense potential forthe health conscious consumers who wish to avoid glucose spikes and thesubsequent initiation of the body's fat storage process due tohigh-glycemic food consumption.

Fabrication of Glucose Sensor Patch

A tattoo-based or patch-based glucose sensor is fabricated on a flexible25 μm thick polyimide, polyester or polyterephthalate sheets using ascreen printer. The substrate is not limited to these materials, andother flexible plastic, textile or elastomeric membranes can also beused. The fabrication comprises of printing a sequence of Ag/AgCl inkand Prussian Blue conductive carbon ink on the flexible substrates usingcustom-designed stencils. The magnetic contact pads (used for attachinga custom-built flexible PCB electronics for wireless communication) arerealized by attaching commercially-obtained conductive and magneticmetal alloy foil by using conductive glue.

FIG. 7 is a schematic of an exemplary tattoo- or patch-basedelectrochemical biosensor device 700 with different components ofmagnetic contacts. The tattoo- or patch-based electrochemical biosensordevice 700 includes a pressure sensitive adhesive membrane 702 and aflexible substrate 704. A cathodic compartment 706, a glucose oxidasefunctionalized electrode 708, an anodic compartment 710, and flexible,conductive, and magnetic contact pads 712 are disposed over the flexiblesubstrate 704. The working electrode is functionalized with the reagentlayer. The enzyme GOx solution (40 mg/mL containing 10 mg/mL BSAstabilizer) is mixed with chitosan solution (0.5 wt % in 1 M aceticacid) in 1:1 v/v ratio. Subsequently, a 2 μL droplet of the abovesolution is casted on the electrode and dried under ambient conditions.Next, 100 μL of 4 wt % agarose gel in 0.1M phosphate buffer is coatedonto the sensor. Finally, the flexible glucose sensor patch is appliedto medical grade pressure sensitive adhesive membrane 702. A wide rangeof commercial urethane, textile based adhesive membranes can be used forthis purpose.

Applying the Glucose Sensor Patch on the Body

The tattoo- or patch-based glucose sensor can be applied to a region ofthe body, such as the deltoid region. A flexible PCB electronics boardis attached to the wearable patch via magnetic attraction between themagnetic contacts pads of the PCB board and that of the glucose patch.FIGS. 8A and 8B are photos of an exemplary glucose sensor patch device800 and exemplary flexible PCB electronics 802. The flexible PCBelectronics 802 provide an effective means to inject current into thebody for the reverse iontophoresis process.

On-Body Studies

The extraction of glucose from the skin interstitial fluid greatlydepends on the contact between the sensor patch and the skin. Conformalcontact between the skin and the sensor leads to efficient glucoseextraction, thus reducing the duration of reverse iontophoresis andpossibility of skin burns. By employing a highly flexible thin polyimidesheet as a substrate for fabricating the glucose sensor along with askin-conforming medical-grade pressure sensitive adhesive, the reverseiontophoresis duration was reduced to 5 min followed by 3 min ofamperometric detection of the extracted glucose as compared to theglucose tattoo sensor platform, that required 10 min of reverseiontophoresis and 5 min of detection. Thus, by using the patch platform,the frequency of glucose detection can be increased while reducing thepossibility of skin irritation. FIG. 9 shows exemplary results for threedifferent subjects obtained by the disclosed glucose sensor patch devicein conjugation with the wireless electronic board. The glucose for eachsubject was measured in the fasting state and after having breakfast.Specifically, FIG. 9 shows amperometric response of glucose sensor patchrecorded for three subjects before and after breakfast A (900), B (902),C (904) and correlation of glucose levels measured by a commercial bloodglucose meter (blue plots) with that measured by the glucose sensorpatch (black plots) for the three subjects A′ (906), B′ (908), C′ (910).

FIG. 10 shows exemplary data from a separate study where the disclosedglucose sensor patch device recorded the glucose spikes occurring aftera subject consumed breakfast, lunch and snacks. For each meal, a newpatch was applied. Specifically, FIG. 10 shows glucose spikes occurringafter consumption of A (1000) breakfast, B (1002) lunch and C (1004)snacks as recorded by glucose sensor patches. D (1006) Correlationbetween glucose levels obtained from a commercial blood glucose meter(blue plot) with that recoded by glucose sensor patches (black plot). Inplot D (1006), the horizontal axis shows BB: before breakfast; AB: afterbreakfast; BL: before lunch; AL: after lunch; BS: before snacks and AS:after snacks.

Tattoo-Based Transdermal Alcohol Sensor

Blood alcohol content (BAC) is most commonly used as an indicator ofalcohol intoxication. However, blood sample is nominally obtained byinvasive means, which gives people pain during sample collection. Toovercome this issue, disclosed is a novel solution to monitor BACnon-invasively in real-time using a wearable tattoo-based biosensornamed ‘AlcoTatt’. Based on a correlation between alcohol concentrationin sweat and blood, the disclosed skin-worn electrochemical biosensormeasures transdermal alcohol content (TAC) in induced sweat viaiontophoresis to estimate BAC, with the measured data wirelesslytransmitted to laptop or mobile device for real-time analysis. The onbody performance of ‘AlcoTatt’ prototype is demonstrated on humansubjects with ingestion of alcohol drinks showing high sensitivity andspecificity towards ethanol in sweat. The present document supports theapplication of a skin-worn tattoo-based wearable electrochemicalbiosensor for the non-invasive alcohol monitoring.

Alcohol consumption leads to harmful consequences such as trafficaccidents and degenerated health care. Therefore, accurate measurementof alcohol consumption is important for preventing alcoholism andalcohol abuse and the effectiveness of their treatment. BAC is mostcommonly used as an indicator of alcohol intoxication. Blood samples areconventionally obtained by invasively pricking a fingers or earlobe,which is a painful process that demands user compliance. The disclosedtechnology provides for an alternative way to measure BAC in anon-invasive, real-time manner.

The disclosed wearable electrochemical sensors can detect metabolitesand electrolytes in a non-invasive way using sweat and interstitialfluid. These sensors can be integrated directly on flexible temporarytattoo substrates for epidermal sensing applications. Suchbody-compliant printable electrochemical sensors provide elasticitycharacteristic of temporary tattoos along with resistance to mechanicalstress and compatibility with the non-planarity of the epidermis.Expanding this attractive skin-worn platform towards non-invasivealcohol detection in sweat benefits wearers comfort and compliance aswell as enables continuous real-time alcohol monitoring, which can beuseful for monitoring clinical treatment status, individuals who areasked to maintain abstinence of alcohol. Also, the device can be used toverify drinking events, prevent driving under the influence, chargingwith driving under influence and track recidivism.

Breathalyzers are the most commonly used device to indirectly estimateBAC by measuring breath alcohol concentration (BrAC). The breathalyzerinstruments calculate BAC by following Henry's law, which hasdifficulties in achieving high accuracy because it can be easilyaffected by humidity, temperature and individuals. BAC also can beestimated by measuring transdermal alcohol concentration (TAC), becausethe person's perspiration can contain traces of alcohol when the personconsumes alcohol. The Giner TAS V is the wearable prototype to measureTAC by detecting the local ethanol vapor concentration over the skin.However, it showed time delay in peak TAS signal compared with BACestimated by breathalyzer varying from 30 min to 2 hours. The disclosedtattoo-based alcohol sensor can monitor BAC non-invasively in real-timeis highly desired. A correlation between ethanol concentration in bloodand sweat during alcohol consumption, with a ratio of 0.81 can beutilized by the disclosed tattoo-based alcohol sensor. The disclosedtattoo-based alcohol sensor is wearable and thus not bulky to carry anddoes not require replacement of electrodes between the iontophoresis andamperometric detection steps.

Specifically, the disclosed wearable tattoo-based alcohol biosensor canbe used to monitor BAC non-invasively in real-time. The disclosedwearable tattoo biosensor provides several distinct advantages andinnovations for practical applications as below. It represents the firstexample of integration of iontophoretic and amperometric-detectionsystem, obviating the need for replacement of electrodes betweeniontophoresis and detection. The electrodes are fabricated byscreen-printing on a wearable tattoo platform, which is easy tofabricate, wear, and remove as well as cost-effective formass-production. Finally, flexible wireless electronics has beenincorporated with tattoo electrodes, which enables real-timenon-invasive measurements and data transmission to lap-top/mobile devicevia Bluetooth communication. The flexible skin-worn board offersresistance to mechanical stress from movement of wearer along withcompatibility with the non-planarity of the epidermis. As described inthis document, the on-body demonstrations reveal that the tattoo-basediontophoretic-biosensing platform holds considerable promise fornon-invasive alcohol monitoring in real-life situations.

FIG. 11 shows an exemplary tattoo-based transdermal alcohol sensordevices and systems. For example, part (A)(i) is a schematic diagram ofiontophoretic-sensing tattoo electrode displaying iontophoreticelectrodes (anode and cathode) and sensing three electrodes (working,reference and counter electrodes) (1100). Part (A)(ii) is a photographof flexible wireless electronics (1102). Part (A)(iii) is a photographof Alcohol iontophoretic sensing tattoo device with integrated flexibleelectronics applied to a human subject (1103). Part (B) is a schematicdiagram of constituents in an exemplary iontophoretic system (1104) andexemplary enzymatic reaction and chemical modification for ethanolsensing on working electrode (1106). Part (C) shows exemplary schematicprocedures of on-body study of alcohol sensing (1108).

An exemplary procedure to monitor alcohol level in sweat using thetattoo-based transdermal alcohol sensor device includes iontophoresisand amperometric detection processes. Iontophoresis is performed toinduce sweat occurrence by delivering pilocarpine as illustrated in FIG.11, part (B) (1104). Enzymetric reaction of alcohol oxidase andmodification of working electrode on Prussian Blue (PB) transducer areshown in FIG. 11, part (B) (1106). In FIG. 11, part (A)(i) (1100), thetattoo electrodes include iontophoretic electrodes (“IE” anode andcathode) and sensing electrodes (CE, WE, and RE: counter, working andreference). For wireless monitoring, the tattoo electrodes can beintegrated with flexible electronics.

As shown in FIG. 11, part (C) (1108), the level of alcohol in sweat canbe monitored under consumption of alcohol beverage by performingionotophoresis (5 minutes), waiting (about 5 minutes), and amperometricdetection by integrated flexible electronics (about 2.5 minutes) anddata can be transmitted wirelessly to a mobile device or lap-top viaBluetooth communication, for example, in real-time. Other wirelesscommunications protocols can also be used.

Fabrication, Chemical Modification, and Transfer Process of TattooSensor

Patterns for printing the sensor were designed in AutoCAD (Autodesk, SanRafael, Calif.) and stainless steel plates (12×12 inch.²) were etched tofabricate stencils (Metal Etch Services, San Marcos, Calif.). Temporarytransfer tattoo paper kits were obtained from HPS Papilio (Rhome, Tex.).A sequence of the silver/silver chloride (Ag/AgCl) ink (E2141 ErconInc., Wareham, Mass.) and prussian blue conductive carbon (Gwent Group,UK) were screen-printed on the substrate by using an MPM-SPMsemi-automatic screen printer (Speedline Technologies, Franklin, Mass.).As illustrated in FIG. 11, the tattoo sensor design includesiontophoresis, pseudo reference, and counter electrodes patterned fromAg/AgCl ink, and the working electrode patterned from Prussian blue ink.A transparent insulator was screen printed over the surface of theelectrode pattern to confine the electrode and contact areas. Eachprinted layer was cured in oven after printing. The Ag/AgCl ink wascured at 90° C. for 10 min, while the prussian blue ink was cured at 80°C. for 10 min in an oven.

In order to obtain ethanol transducer, the working electrode was thenfunctionalized with the enzymatic layer (BSA, chitosan, and AOx enzyme).The AOx enzyme, BSA stabilizer and chitosan solution (0.5 wt % in aceticacid solution) was mixed together in an 8:1:1 v/v ratio. Afterward, a 4μL droplet of the mixed solution was casted on the electrode. Afterair-drying, the electrode was covered with a 2 μL chitosan solution.Then, the electrode was dried under ambient conditions. The agarosehydrogel was prepared by heating a continuously stirred agarose solution(4% w/v) in 0.1 M potassium phosphate buffer (pH 7.0) until agarose iscompletely dissolved. Dissolved agarose hydrogel is then casted on theworking electrode. Then prepared PVA cryogels (2.0×1.5 cm²) were soakedin 1% pilocapine nitrate and 1% sodium nitrate and then covered on anodeand cathode compartments, respectively. The PVA cryogels were made asfollowing procedure by freezing and thawing sequences of PVA solution.First, 5.0% v/v PVA solution was prepared in deionized water by heatingthe solution to 120° C. and then cooled to room temperature. After themixture is cooled, it was placed in the ice bath and pH was adjusted topH 1 by adding 5 M hydrochloric acid. Subsequently, the cross-linkerglutaraldehyde was added to make a final concentration 0.5% w/v. Thefinal mixture was then stirred for 1 minute and poured onto the petridish to set in −20° C. freezer overnight. The cured PVA cryogel waspatterned by paper cutter to control size of applying iontotophoresis.

Evaluation of Sensor Performance in Buffer Medium

The electrochemical performance of the alcohol tattoo biosensor wasfirst tested in 0.1M phosphate buffer (pH 7.0) medium.Chronoamperometric response was measured by stepping the potential to−0.2V (vs. Ag/AgCl) for 60 s after 1 min incubation. The calibrationcurve was obtained with 3 mM increments of ethanol concentrations up to36 mM in buffer solution. Selectivity was examined by response of 10 mMethanol in the presence of relevant electroactive species: 0.2 mMglucose, 10 mM lactate, 84 μM creatine, 10 μM ascorbic acid, and 60 μMuric acid.

Evaluate the On-Body Performance of Wearable Tattoo Alcohol Biosensor

The epidermal evaluation on human subjects was conducted in strictcompliance following a protocol approved by the institutional reviewboard (IRB) at the University of California, San Diego. Total ninehealthy volunteers were recruited for on-body evaluation of thedeveloped sensor under taking alcohol beverages. First, tattoo biosensorwas transferred on subject's arm and a set of ethanol detection in sweatis followed to get a current response at BAC 0.00% (‘before drinking’).Experimental time frame for on-body experiments is illustrated in FIG.1C. The set of ethanol detection consists of iontophoresis andamperometry. During iontophoresis process, a constant current of 0.6 mA(0.2 mA/cm²) was applied through PVA cryogel between the twoiontophoresis electrode (anode and cathode) for 5 min to deliverpilocarpine chemical towards skin to induce sweat. After iontophoresis,5 min of resting was required to give time to generate sweat. Finally,amperometric response of ethanol in sweat was recorded at appliedpotential of −0.2 V (vs Ag/AgCl) for 150 sec, corresponding to thecurrent response at BAC 0.00%. The subject was asked to take analcoholic beverage (12 oz. of beer or 5 oz. of table wine) and waitedfor 10 min to let alcohol diffuse in blood stream. One set ofiontophoresis/detection cycle was followed to check the response ofethanol in sweat. Along with on-body experiments taking alcoholbeverages, three different types of control experiments (withoutdrinking, without enzyme modification, without iontophoresis) wereperformed. The sensing response towards sweat ethanol was confirmed bytesting without drinking alcoholic beverages.

Furthermore, additional on-body experiments are followed to verifycorrelation between BAC and current response from our tattoo sensor withserial consumption of wine drinks. As described in previous paragraph,same experimental procedure was conducted by repeating two sets of drinkand measurement. Before drinking (when BAC is 0.00%), current responsewas recorded. Then, after 1^(st) drinking, current response was obtainedand followed by 2^(nd) drinking and measurement. Each set of measurementcycle is accompanied with simultaneous measurement of BAC usingcommercial breathalyzer (Alcovisor Mars breathalyzer) to validate oursensor performance in comparison with commercial alcohol sensor.

Design and Fabrication of Wireless Flexible Printed Circuit Board

This work introduces a complete flexible wearable device fornon-invasive monitoring of alcohol (FIG. 1). The system includes atattoo-based iontophoresis alcohol monitoring system, as well as aflexible wireless electronic system that transmits the alcohol levelinformation through a Bluetooth low energy (BLE) link (FIGS. 1, A andB). A BLE-enabled printed circuit board has been designed to implement aprototype for the iontophoresis alcohol biosensor system. The circuitemploys a Texas Instrument (TI) CC2541 BLE System-on-Chip forcommunication and processing. A Texas Instrument LM334, current source,applies 0.6 mA current between cathode and anode electrodes. A TexasInstrument LMP91000 (analog front end for chemical sensing) was used asthe amperometric monitoring system for the alcohol sensor. A current of600 uA is applied between the anode and cathode electrodes for 5minutes. Afterwards, the CC2541 microcontroller disconnects the currentsource and activates the amperometric chip to measure the amperometriccurrent for a potential cell of −0.2 V after 5 min delay. Then, thesensor information is transmitted in a 2-byte format to a Bluetooth4.0-enabled reciever. A graphical interface has been developed usingPython script language to demonstrate measurement results on a desktopor laptop. A Johanson Technology 2.45 GHz chip antenna (2450AT42A100)and impedance matched balun (2450BM15A0002) were employed for wirelesstransmission. Two 396/397 watch batteries (2×1.55 V, 33 mAh each) inseries were utilized as a power source, regulated for the electronicsvia a TPS61220 boost converter and an LM4120 low-dropout voltageregulator.

Assembly and Characterization of Integrated Wireless Tattoo-BasedAlcohol Sensor

The first prototype of the fabricated flexible printed circuit boardassembly, shown in FIG. 1A, measured 2 cm×5 cm. In the flexible PCBdesign, the holes were selectively plated inside so that the surfacecopper stays at 0.5 oz RA cu with no plated copper on top. That is ineffort to minimize the possibility of cracking the traces whilemishandling or bending the flex. Unlike the usual process, in whichPolyimide cover layer is used for insulation, due to the complexity ofthe soldermask openings, LPI is employed to get a good registration.Then, Polyimide cover layer provides extra support for the surfacecopper when bending.

Rationale for Iontophoretic-Biosensing System of Tattoo-Based AlcoholSensor

Each pair of alcohol tattoo sensor consists of iontophoretic electrode(anode and cathode) and amperometric sensing electrode (WE, RE, and CE)in anode compartment. Due to integration of both systems in oneplatform, specific electrode design was required. The iontophoreticelectrodes are responsible for generating sweat by delivery ofpilocarpine drug and thus positioned in the middle of sensing electrodecomponent. Iontophoresis electrodes were covered with cryogel soakedwith pilocarpine nitrate and sodium nitrate on anode and cathode,respectively, to deliver the chemicals though skin by applying constantcurrent. The cryogel has large porous structure and biocompatiblematerial offering effective sweat generation without any skin irritationand burning in our preliminary study. For amperometric sensing ofalcohol in sweat, screen-printed Prussian-Blue transducer was utilizeddue to its high selectivity towards hydrogen peroxide, which is productof enzymatic reaction between alcohol oxidase (AOx) and alcohol. AOx wasimmobilized on working electrode with BSA, chitosan and then coveredwith agarose gel containing PBS (K⁺) to provide enough electrolytes torun electrochemistry, especially potassium ion which is crucial ion tokeep electron shuttling activity of Prussian-Blue.

Determination of Alcohol with Tattoo Sensor in Buffer Medium

First of all, the electrochemical performance of the alcohol biosensorwas validated in buffer medium over a dynamic concentration range of0-36 mM ethanol, which is physiological level in sweat. FIG. 12, part Adisplays well-defined chronoampemetric response to 3 mM increment (b-m)in buffer solution (1200). The resulting linear calibration plot isshown in the inset of FIG. 12, part A (slope, 0.441 μA/mM; correlationcoefficient, R²=0.992). Note that human sweat contains variousphysiological relevant interferents (glucose, uric acid, lactate,ascorbic acid, and creatine), selectivity towards ethanol should betested for on-body operation and the result is shown in FIG. 12, part B(1202). High sensitivity and selectivity toward ethanol are demonstratedin FIG. 12 and these are attributed by efficient chemical modificationof alcohol oxidase on Prussian blue working electrode and highselectivity of Prussian blue transducer towards hydrogen peroxide(product of enzymatic reaction) at low operational potential.

FIG. 12 shows exemplary results obtained using the disclosednon-invasive biosensor device implemented as a tattoo- or patch-basedalcohol sensor. Specifically, FIG. 12 shows in part (A),chronoamperometric response of the tattoo-based alcohol sensor toincreasing ethanol concentrations from 0 mM(a) to 36 mM(m) in buffer in3 mM increments (1200). FIG. 12 shows in part (B), interference study inthe presence of 10 mM ethanol (plot “a”), followed by subsequentadditions of 0.2 mM glucose (plot “b”), 10 mM lactate (plot “c”), 84 μMcreatine (plot“d”), 10 μM ascorbic acid (plot “e”), and 60 μM uric acid(plot “e”. Potential step to −0.2 V (vs Ag/AgCl). Medium,phosphate-buffer (pH 7) (1202). FIG. 12 shows in part (C), cyclicvoltammogram of the tattoo-based alcohol sensor in buffer solution (pH7) and 1% pilocarpine dissolved buffer solution (pH 7) (1204). FIG. 12shows in part (D), cyclic voltammogram of the tattoo-based bare prussianblue (PB) electrode and 2% agarose gel modified PB electrode in 1%pilocarpine solution (dissolved in DI water) (1206).

On-Body Alcohol Monitoring on Human Subject

After evaluation of the tattoo sensor performance in vitro, we testedon-body operation of alcohol tattoo sensor with human subjects. First,detection ability of the alcohol tattoo sensor is confirmed underconsumption of alcoholic beverage. FIG. 13 A shows on-body resultsobtained by three different subjects with alcohol ingestion followingthe protocol described in experimental approach section (1300, 1302,1304). Amperometric responses were compared between before and afterdrinking alcohols, showing distinct current signal increment caused byspike of alcohol level in sweat along with its level in blood over allthree subjects. As described, second amperometric response (afterdrinking) was measured after 20 min of alcohol consumption and BAC ismeasured at the same time with amperometric measurement. Even thoughthree subjects were asked to take same amount of alcohols, the BACvalues and current responses are different among subjects after thegiven time of period, implying each subject has different digestion rateof alcohol. None of the subjects reported perceptible discomfort duringthese on-body measurements. Three control experiments (no drinking, noenzyme modification, and no iontophoresis) were performed to verifyfalse alarm from on-body results obtained under alcohol consumption.None of control experiments showed current difference. In detail, thedata from control experiment without enzyme modification clearlydemonstrates high specificity of alcohol tattoo sensor toward alcohol insweat. It can be clearly noted from control experiment without drinking(FIG. 13B (1306, 1308, 1310)), that the current response, shown in FIG.13A (1300, 1302, 1304), is caused by alcohol consumption and performingiontophoresis process doesn't affect any amperometric current signal,even though remained amount of pilocarpine nitrate has changed due todelivery of pilocarpine. The last control experiments in the absence ofiontophoresis also showed no current differences although BAC was spikedup to 0.0018% from alcohol ingestion. It implies that current responsesshown in FIG. 13A (1300, 1302, 1304) are from alcohol in generated sweatby iontophoresis not from other source/environment.

Specifically, FIGS. 13A-B represents amperograms obtained using thedisclosed noninvasive alcohol sensing from human subjects wearing thedisclosed alcohol tattoo sensor. Specifically, FIG. 13 (A) showsexperiments with consumption of 12 oz. of beer measured before and afterdrinking alcohol beverage from three subjects (1300, 1302, and 1304).FIG. 13 (B) shows control experiments without drinking (left 1306),without enzyme immobilization (middle 1308), and without iontophoresis(right 1310). Blood alcohol level is obtained by a breathalyzer. In FIG.13, data were collected under potential step to −0.2 V (vs Ag/AgCl).

Next, additional on-body experiments are performed to evaluatecorrelation between BAC (obtained from breathalyzer) and currentresponse from our alcohol tattoo sensor with serial ingestion of alcoholdrinks. As illustrated in FIG. 13C, first current response was obtainedcorresponds to BAC 0.00% (before drinking) and followed by two sets ofdrinking and measurement procedures using same electrode (1320). After1^(st) alcohol consumption, BAC turned out 0.025% showing distinctcurrent response, and BAC was increased to 0.062% after 2^(nd) drinkingwith further increased amperometric current signal. The resulting plotbetween BAC and current value is shown in FIG. 13D (1322). The currentresponse displays great linearity toward BAC level (slope, 102 μA/BAC %;correlation coefficient, R²=0.999). This implies that the currentresponse is responsible for ethanol level in sweat, which iswell-correlated with BAC.

Specifically, FIG. 13C shows exemplary Amperograms obtained fornoninvasive alcohol detection obtained from human subjects wearing thealcohol tattoo sensor showing correlation between BAC level and currentresponse from tattoo biosensor measured before (i, BAC: 0%) and afterdrinking 5 oz. of wine (ii, BAC: 0.0025%)) and 10 oz. of wine (iii, BAC:0.0062%)). Blood alcohol level is obtained by breathalyzer. Potentialstep to −0.2 V (vs Ag/AgCl). FIG. 13D shows resulting linear plotbetween current response and BAC level.

Characterization of Integrated Flexible Wireless Tattoo-Based AlcoholSensor

The practical use of wearable biosensors for real-time monitoring hasbeen hindered by the lack of the development of a body-compliantwireless circuit board. For realization of wearable alcohol sensor, aflexible wireless circuitry was custom-made and integrated with ourdeveloped sensor (shown in FIG. 11, part A). Performance of theintegrated wearable device was evaluated following the same protocol asprevious on-body experiments. A BLE-enabled printed circuit boardemployed iontophoresis by applying 0.6 mA current and measuredamperometry at −0.2 V for 150 seconds. The current response was sampledwith a frequency of 1 Hz, and transmitted in real time via Bluetooth 4.0to a laptop/mobile device and plotted on screen with graphical interfacedeveloped using Python. The resulting plot is shown in FIG. 13E obtainedby wireless transmission showing clear current difference afterconsumption of alcohol beverage and the data corresponds to previouson-body results obtained from lab-scale potentiostat. This test providesthe evidence for the practicality of the wireless alcohol tattoo sensortowards monitoring BAC.

Specifically, FIG. 13E shows integration of flexible wireless circuitboard with alcohol tattoo sensor. Amperograms obtained by the wirelessintegrated alcohol tattoo sensor from human subjects measured before andafter taking alcohol beverage. BAC was confirmed as 0.035% bybreathalyzer. Potential step to −0.2 V (vs Ag/AgCl).

In this patent document, ‘AlcoTatt’, skin-worn tattoo-based wearableelectrochemical biosensors are disclosed for noninvasive alcoholmonitoring. Based on well-established correlation between ethanolconcentration in sweat and blood, AlcoTatt measures alcoholconcentration in sweat to estimate BAC, allowing non-invasive monitoringof BAC in real-time. The in vitro characterization of the tattoo sensorsrevealed their ability to detect alcohol in sweat covering physiologicalrange in the presence of common interfering chemical species. On-bodyevaluation of the tattoo-based iontophoretic-biosensing platform furtherdemonstrated the ability to detect the rise in the ethanol level afterconsumption of alcohol beverage in a noninvasive fashion. The new tattooalcohol sensor has been coupled with flexible printed circuit board forwireless data collection in real-time. The disclosed ‘AlcoTatt’ allowsnon-invasive, passive and simple monitoring of BAC, which can be usefulfor checking driver's ethanol ingestion or individuals who need to keepabstinence of alcohol.

Alcohol Tattoo Sensor Designs

The disclosed technology can be implemented as alcohol tattoo sensordevices. FIG. 14 shows an exemplary Pilocarpine Induced Sweat Generationusing an exemplary alcohol tattoo sensor device 1400. In FIG. 14, thePilocarpine Induced Sweat Generation was performed using a PilocarpineSoaked in Filter Paper. Table 1 shows exemplary amounts of sweatgenerated during 10 min collection time.

TABLE 1 Amount of sweat generated during 10 min collection time Before(mg) After (mg) Difference Un-stimulated 32.94 33.5 0.56 Stimulated31.45 82.55 51.6

FIG. 15 shows an exemplary on-tattoo electrode 1500. An iontophoresiselectrode is used as counter electrode 1502. The on-tattoo electrode1500 also includes a reference electrode 1504 and a working electrode1506. When the electrode 1500 is oxidized, the color changes to black,for example. The on-tattoo was tested at −0.2V with 1 min incubation toobtain an EtOH response. The on-tattoo can be optimized in terms ofpotential, and enzyme amount.

FIG. 16 shows exemplary data on alumina SPE_EtOH detection (1600). Thedata 1600 was obtained for Chitosan (1602) and Nafion (1604).

FIG. 17 shows exemplary data 1700 and 1702 on EtOH detection on Aluminasubstrate (PG design). The data was obtained under electrodemodification that includes 1 mL of (2.5 mL of (AOx 20 U/mL+BSA 10mg/mL)+0.5 mL Chitosan 0.5% v:v). Working and counter electrodes includePB carbon material. The reference electrode includes Ag/AgCl.

FIG. 18 shows exemplary data 1800 and 1802 on EtOH detection on Tattoopaper. The working electrode includes PB carbon while the reference andcounter electrodes include Ag/AgCl. At the working electrode, reductioncurrent occurs, at the same time, silver counter electrode is easilyoxidized.

FIG. 19 shows data for exemplary EtOH detection with Std.3 design ontattoo (PG stencil) 1900 and 1902. The working and counter electrodeswere PB carbon and only the reference electrode was Ag/AgCl. The data1900 and 1902 were obtained under the working electrode modification of1 mL of (2.5 mL of (AOx 20 U/mL+BSA 10 mg/mL)+0.5 mL Chitosan 0.5% v:v).Troubleshooting can be performed by changing the Counter Electrode to PBcarbon instead of Ag/AgCl. Consequently, the Current signal is in abetter shape showing linear response to ethanol concentration.

FIG. 20 shows data 2000, 2002, and 2004 for exemplary EtOH detectionwith Std.3 design on tattoo (PG stencil). The working electrode and thecounter electrodes were PB carbon and only the reference electrode wasAg/AgCl. The data 2000, 2002, and 2004 were obtained with workingelectrode modification of (Aox+BSA)+0.5% Chitosan:double-layer. Foramperometry, applied potential was −0.2V. The data recording protocolincluded 0.1 M phosphate buffer (pH 7.4) and 50 mV/s, 2nd cycle.

The blood alcohol content (BAC) can be calculated based on the sweatethanol concentration using Equation 1. For example, 0.01% BAC=0.04 mMEtOH in sweat and 0.08% BAC=0.35 mM EtOH in sweat.

BAC (g L⁻¹)=0.71×sweat ethanol concentration (g L¹) (r=0.9).  Eq. 1

According to a DMV BAC regulation, 0.01% BAC (0.04 mM EtOH) should bedetectable. To perform alcohol sensing, carbon counter can be usedrather than Ag/AgCl. EtOH affects tattoo paper substrate, and theamperometric response can be unstable due to bad connection. Additionalinsulator layer can be printed on tattoo paper, or on PET substrate toaddress the issue.

FIG. 21 shows an exemplary tattoo substrate 2100. Ethanol may damage thetattoo paper layer. Troubleshooting can be performed by transferring thetattoo electrode on petri dish (or plastic substrate) and consequentlythe electrode is working fine.

FIG. 22 shows exemplary electrode design and system for 2E and 3Esystems 2200. An exemplar glucose sensing tattoo device design caninclude 1 silver iontophoresis electrode (Big)+2E-system for Glucosedetection. The ability of the 2E system to detect ethanol effectivelycan be affected when the reference electrode, counter electrode, or bothare oxidized. When the iontophoresis electrode is used as the counterelectrode, the silver counter electrode may still be oxidized because itis easily oxidized. The 3E system with Prussian Blue carbon electrode ismore effective for Ethanol system because the Prussian Blue carbonelectrode is not easily oxidized. Potential solutions can include (1)adding a carbon counter electrode on the glucose tattoo design; (2)adding an iontophoresis electrode (Silver) on PG stencil; or (3) makinga new design for ethanol sensor.

FIG. 23 shows data 2300 and 2302 from an exemplary glucose tattooelectrode with 3E system (Silver counter and ref electrode.) The data2300 and 2302 were obtained under enzyme modification ofAox+BSA+Chitosan (1 layer) E=−0.2V incubation 1 min.

FIG. 24 shows data 2400 and 2402 obtained from an exemplary PG tattooElectrode with 3E system (Prussian Blue carbon electrode as the counterelectrode). The data 2400 and 2402 were obtained under enzymemodification of Aox+BSA+Chitosan (1 layer) E=−0.2V incubation 1 min.FIG. 24 shows the response to 1 mM.

As shown above, using the proper tattoo design for Ethanol detection iscrucial. In addition, enzyme modification can be optimized to obtain agood linear range in physiological level.

FIG. 25 are schematic diagrams and real images displaying an exemplarybiodevice (2500, 2504) and enzyme and electrode reactions involved inthe response of graphite-Teflon-AOD-HRP-ferrocene (2502, 2506).

FIG. 26 shows an exemplary amperometric trace 2600 measured with anexemplary biodevice at 0.00V depicted for a 5.0×10⁻⁴ molL⁻¹ ethanolsolution.

FIG. 27 show exemplary current-time recording obtained after alcoholingestion by placing a biodevice on the skin: continuous mode recordingfor three different volunteers (2700) and single measurements carriedout for one volunteer (2702). In FIG. 27, the E_(app)=0.00 v.

FIG. 28 shows exemplary correlation data 2800 between the cathodiccurrent values measured with an exemplary biodevice in the singlemeasurement mode for 40 volunteers (filled square) as well as others(open circle) vs. the BAC values determined by gas chromatography.Calibrating straight line prediction intervals are shown as solid lines.

FIG. 29 shows least square straight line regression data 2900 resultingby plotting the BAC values in gL⁻¹ obtained with the biodevice measuringin sweat and the gas chromatography reference method.

FIG. 30 shows a comparison data 3000 of the BAC values obtained with thebiodevice through sweat measurements and with the gas chromatographymethod.

Implementations of the subject matter and the functional operationsdescribed in this patent document can be implemented in various systems,digital electronic circuitry, or in computer software, firmware, orhardware, including the structures disclosed in this specification andtheir structural equivalents, or in combinations of one or more of them.Implementations of the subject matter described in this specificationcan be implemented as one or more computer program products, i.e., oneor more modules of computer program instructions encoded on a tangibleand non-transitory computer readable medium for execution by, or tocontrol the operation of, data processing apparatus. The computerreadable medium can be a machine-readable storage device, amachine-readable storage substrate, a memory device, a composition ofmatter effecting a machine-readable propagated signal, or a combinationof one or more of them. The term “data processing apparatus” encompassesall apparatus, devices, and machines for processing data, including byway of example a programmable processor, a computer, or multipleprocessors or computers. The apparatus can include, in addition tohardware, code that creates an execution environment for the computerprogram in question, e.g., code that constitutes processor firmware, aprotocol stack, a database management system, an operating system, or acombination of one or more of them.

A computer program (also known as a program, software, softwareapplication, script, or code) can be written in any form of programminglanguage, including compiled or interpreted languages, and it can bedeployed in any form, including as a stand-alone program or as a module,component, subroutine, or other unit suitable for use in a computingenvironment. A computer program does not necessarily correspond to afile in a file system. A program can be stored in a portion of a filethat holds other programs or data (e.g., one or more scripts stored in amarkup language document), in a single file dedicated to the program inquestion, or in multiple coordinated files (e.g., files that store oneor more modules, sub programs, or portions of code). A computer programcan be deployed to be executed on one computer or on multiple computersthat are located at one site or distributed across multiple sites andinterconnected by a communication network.

The processes and logic flows described in this specification can beperformed by one or more programmable processors executing one or morecomputer programs to perform functions by operating on input data andgenerating output. The processes and logic flows can also be performedby, and apparatus can also be implemented as, special purpose logiccircuitry, e.g., an FPGA (field programmable gate array) or an ASIC(application specific integrated circuit).

Processors suitable for the execution of a computer program include, byway of example, both general and special purpose microprocessors, andany one or more processors of any kind of digital computer. Generally, aprocessor will receive instructions and data from a read only memory ora random access memory or both. The essential elements of a computer area processor for performing instructions and one or more memory devicesfor storing instructions and data. Generally, a computer will alsoinclude, or be operatively coupled to receive data from or transfer datato, or both, one or more mass storage devices for storing data, e.g.,magnetic, magneto optical disks, or optical disks. However, a computerneed not have such devices. Computer readable media suitable for storingcomputer program instructions and data include all forms of nonvolatilememory, media and memory devices, including by way of examplesemiconductor memory devices, e.g., EPROM, EEPROM, and flash memorydevices. The processor and the memory can be supplemented by, orincorporated in, special purpose logic circuitry.

Additional information pertaining to the disclosed technology isdescribed below.

Wearable Electrochemical Sensors

In some implementations, the disclosed technology can be applied towearable sensors. Sensors based on electrochemical processes can be usedto detect a chemical, substance, a biological substance (e.g., anorganism) by using a transducing element to convert a detection eventinto a signal for processing and/or display. Biosensors can usebiological materials as the biologically sensitive component, e.g., suchas biomolecules including enzymes, antibodies, nucleic acids, etc., aswell as living cells. For example, molecular biosensors can beconfigured to use specific chemical properties or molecular recognitionmechanisms to identify target agents. Biosensors can use the transducerelement to transform a signal resulting from the detection of an analyteby the biologically sensitive component into a different signal that canbe addressed by optical, electronic or other means. For example, thetransduction mechanisms can include physicochemical, electrochemical,optical, piezoelectric, as well as other transduction means.

Techniques, systems, and devices are described for fabricating andimplementing electrochemical biosensors and chemical sensors that arewearable on skin or a wearable item, e.g., by a procedure analogous tothe transfer of a temporary tattoo.

In one aspect of the disclosed technology, a method of producing anepidermal biosensor includes forming an electrode pattern onto a coatedsurface of a paper-based substrate to form an electrochemical sensor,the electrode pattern including an electrically conductive material andan electrically insulative material configured in a particular designlayout, and attaching an adhesive sheet on a surface of theelectrochemical sensor having the electrode pattern, the adhesive sheetcapable of adhering to skin or a wearable item, in which theelectrochemical sensor, when attached to the skin or the wearable item,is operable to detect chemical analytes within an external environment.

Implementations of the method can optionally include one or more of thefollowing features. For example, in some implementations of the method,the adhesive sheet can include an outer coating layer on an externalsurface of the adhesive sheet not in contact with the electrode pattern.For example, the outer coating layer can include polyvinyl alcohol(PVA). In some implementations, for example, the method can furtherinclude removing the outer coating layer from the adhesive sheet toenable adhesion of the electrochemical sensor to the skin or thewearable item via the adhesive sheet. Also for example, the method canfurther include removing the paper-based substrate from theelectrochemical sensor to expose the electrode pattern to the externalenvironment. For example, the coated surface can include a release agentmaterial including cellulose acetate. For example, the adhesive sheetcan include polydimethylsiloxane (PDMS). In some implementations of themethod, the forming can include performing screen printing, aerosoldeposition, or inkjet printing the electrode pattern onto the coatedsurface of the paper-based substrate. For example, the electricallyconductive material can include a conductive ink, e.g., including, butnot limited to, gold, platinum, nickel, copper, silver, and/or silverchloride. For example, the electrically insulative material can includea nonconductive ink, e.g., including, but not limited to, polyethyleneterephthalate (PET), polystyrene (PS), polyester (PE), and/orpolytetrafluoroethylene (PTFE). In some examples, the electrode patterncan further include an electrically semi-conductive material. Forexample, the electrically semi-conductive material can include asemi-conductive ink, e.g., including, but not limited to, amorphouscarbon, carbon black, graphite, carbon nanotubes, and/or graphene. Insome implementations of the method, for example, the electrode patternfurther can include carbon fiber segments dispersed within theelectrically conductive or electrically semi-conductive material.

In another aspect, a method of producing an epidermal biosensor includesforming an electrode pattern onto a coated surface of a paper-basedsubstrate to form an electrochemical sensor, the electrode patternincluding an electrically conductive material and an electricallyinsulative material configured in a particular design layout, attachingan adhesive sheet on a surface of the electrochemical sensor having theelectrode pattern, the adhesive sheet capable of adhering to skin or awearable item and structured to include a coating layer on an externalsurface of the adhesive sheet, and removing the paper-based substratefrom the electrochemical sensor to expose the electrode pattern, inwhich the electrochemical sensor, when attached to the skin or thewearable item, is operable to detect a substance present within a fluidthat contact the electrode pattern coupled to the skin or the wearableitem.

Implementations of the method can optionally include one or more of thefollowing features. For example, in some implementations of the method,the electrochemical sensor can be operable to detect physiological,biological, or chemical signals from the skin. In some implementations,for example, the method can further include, when attached to the skinor the wearable item, removing the coating layer from the adhesive sheetexposing a non-adhesive surface of the adhesive sheet. For example, thecoating layer can include PVA. For example, the coated surface of thepaper-based substrate can include a release agent material includingcellulose acetate. For example, the adhesive sheet can include PDMS. Insome implementations of the method, the forming can include performingscreen printing, aerosol deposition, or inkjet printing the electrodepattern onto the coated surface of the paper-based substrate. Forexample, the electrically conductive material can include a conductiveink, e.g., including, but not limited to, gold, platinum, nickel,copper, silver, and/or silver chloride. For example, the electricallyinsulative material can include a nonconductive ink, e.g., including,but not limited to, PET, PS, PE, and/or PTFE. In some examples, theelectrode pattern can further include an electrically semi-conductivematerial. For example, the electrically semi-conductive material caninclude a semi-conductive ink, e.g., including, but not limited to,amorphous carbon, carbon black, graphite, carbon nanotubes, and/orgraphene. In some implementations of the method, for example, theelectrode pattern further can include carbon fiber segments dispersedwithin the electrically conductive or electrically semi-conductivematerial.

In another aspect, an epidermal electrochemical sensor device includes asubstrate formed of a flexible electrically insulative materialstructured to adhere to skin or a wearable item, a first electrodeformed on the substrate of an electrically conductive material, a secondelectrode configured on the substrate of a material that is electricallyconductive and separated from the first electrode by a spacing region,the first and second electrodes capable of sustaining a redox reactionto produce an electrical signal, and a first electrode interfacecomponent and second electrode interface component formed on thesubstrate and electrically coupled to the first electrode and the secondelectrode, respectively, via electrically conductive conduit, in which,when attached to the skin or the wearable item and electrically coupledvia the first and second electrode interface components to one or moreelectrical circuits, the device is operable to detect a substance in alocal environment of the skin or the wearable item.

Implementations of the device can optionally include one or more of thefollowing features. For example, in some implementations of the device,at least one of the first electrode or the second electrode can includean enzyme catalyst and an electroactive redox mediator, theelectroactive redox mediator facilitating the transfer of electronsbetween the electrode and the active site of the enzyme catalystconfigured to sustain a redox reaction. In some implementations, forexample, the device can further include an electrically conductiveunderlayer on the substrate and underneath each of the first electrodeand the second electrode, respectively, the underlayer providingseparation of the first electrode and the second electrode.

In another aspect, a method to fabricate an epidermal electrochemicalsensor device includes depositing an electrically conductive ink on anelectrically insulative substrate to form two or more electrodesadjacent to and separated from one another and conduit wires connectingto each of the electrodes, the depositing including printing the ink ona first stencil placed over the substrate, the first stencil including apatterned region configured in a design of the two or more electrodesand the conduit wires to allow transfer of the ink on the substrate, andthe first stencil inhibiting transfer of the ink in areas outside thepatterned region; curing the electrically conductive ink; depositing anelectrically insulative ink on the substrate to form an insulative layerthat exposes the two or more electrodes, the depositing includingprinting the electrically insulative ink on a second stencil placed overthe substrate, the second stencil including a printing region configuredin a second design to allow transfer of the ink on the substrate, thesecond stencil inhibiting transfer of the ink in areas outside theprinting region; and curing the electrically insulative ink.

Implementations of the method can optionally include one or more of thefollowing features. For example, in some implementations, the method canfurther include depositing an adhesive layer on the insulative layerthat exposes the two or more electrodes, the adhesive substrate formedof a flexible electrically insulative material structured to adhere toskin or a wearable item of a user. In some implementations, for example,the substrate can include a paper substrate. For example, the papersubstrate can include an upper layer and a base paper layer, the upperlayer comprising a release agent coated on the base paper layer andstructured to peel off to remove the paper substrate. For example, thecuring can include implementing at least one of applying heat orultraviolet radiation to the deposited ink on the substrate. In someimplementations, for example, the method can further include forming anelectrically semi-conductive layer over at least one of the two or moreelectrodes by printing an ink of an electrically semi-conductivematerial on a third stencil placed over the substrate, the third stencilincluding a printing region configured in a first design of the at leastone of the two or more electrodes, the printing region allowing transferof the ink on the paper substrate, and the third stencil inhibitingtransfer of the ink in areas outside the printing region; and curing theelectrically semi-conductive ink. In some implementations, for example,the method can further include dispersing carbon fibers in theelectrically conductive ink. In some implementations, for example, themethod can further include depositing an ion-selective membrane to thesurface of at least one of the electrodes, in which the depositingincludes performing at least one of: (i) drop-casting the ion-selectivemembrane on the anterior surface of the electrode, (ii) screen printingthe ion-selective membrane on the anterior surface of the electrode,(iii) inkjet printing the ion-selective membrane on the anterior surfaceof the electrode, and/or (iv) aerosol deposition of the ion-selectivemembrane on the anterior surface of the electrode. In someimplementations, for example, the method can further include depositinga catalyst to the surface of at least one of the electrodes, in whichthe depositing includes performing at least one of: (i) encasing thecatalyst in a porous scaffold structure formed of a conducting polymeron the surface of the electrode, (ii) covalently binding the catalyst tothe surface of the electrode, (iii) entrapping the catalyst in aselectively permeable membrane coupled to the surface of the electrode,and/or (iv) electrostatically binding the catalyst to the surface of theelectrode. In some implementations, for example, the method can furtherinclude depositing an electroactive redox mediator to the surface of theat least one of the electrodes including the catalyst to form anelectrochemical sensing layer, in which the electroactive redox mediatorfacilitates the transfer of electrons between the electrode and theactive site of the catalyst. In some implementations, for example, themethod can further include depositing multi-walled carbon nanotubes onthe surface of at least one of the two or more electrodes.

The subject matter described in this patent document can be implementedin specific ways that provide one or more of the following features. Forexample, the disclosed technology has wide-ranging implications in thehealthcare, fitness, sport and athletics performance monitoring,beauty/skin care, dermatology, environmental, and general sensingdomains. For example, the disclosed technology can be easily adapted foruse in the generalized healthcare, fitness, sport, remote monitoring,wireless healthcare, personalized medicine, performance monitoring, andwar-fighter monitoring domains. Also, for example, the disclosedtechnology can involve the substitution of test strips for metaboliteand electrolyte monitoring in the perspiration, and may replaceconventional screen printed electrochemical test strips in otherdiagnostics and environmental monitoring applications.

Advances in material and device fabrication techniques can be usedbody-worn electronic devices that are mated directly with the skin forthe measurement of physiological parameters of the individual wearer.The disclosed technology here can enable body-worn devices to provideanalysis of chemical constituents residing on the surface of the skin.This analysis can provide useful insight into the overall health andphysical activities of an individual and possible exposure to chemicalor biological agents and certain hazardous substances. For successfulimplementation of direct epidermal electrochemical devices attached tothe body of a user, the devices must exhibit compatible elasticitybetween the device substrate and the skin.

Techniques, systems, and devices are described for fabricating andimplementing electrochemical biosensors and chemical sensors that can betransferred onto the skin or wearable item, e.g., by a procedureanalogous to the transfer of a temporary tattoo.

The disclosed technology can include biosensors and chemical sensorsthat use detection methodologies including amperometry, voltammetry,potentiometry, and/or electrochemical impedance spectroscopy forepidermal monitoring of a wearer's bodily substances such as fluids orgases or the wearer's exposure to one or more substances in thesurrounding environment. The disclosed biosensors and chemical sensorscan include temporary transfer electrochemical biosensors and chemicalsensors that can be applied for limited or long term use on the user'sskin for direct physiological and security monitoring of chemicalconstituents. The exemplary temporary transfer electrochemicalbiosensors and chemical sensors can be produced in the form of aestheticdesigns similar as a skin tattoo, referred to herein as epidermaltemporary transfer tattoo (T3) sensors. In some implementations, thedisclosed sensors can include electrode patterns forming a completeelectrochemical system, as well as include the selection of theappropriate layering and ink formulation to facilitate the requiredelectrochemical response. Exemplary sensors of the disclosed technologycan be implemented in epidermal monitoring of the wearer's environmentand physiological fluids residing on the surface of the epidermis, e.g.,such as perspiration. For example, the exemplary sensors can be used tomeasure one or more physiological parameters, e.g., including but notlimited to measurements of chemical or biological substances in bodyfluids and provide useful insight into the real-time physical conditionsor overall health of the individual wearer as well as their exposure tochemical or biological agents/hazards residing in their localenvironment by analyzing of the detected chemical constituents residingon the surface of the skin. For example, the exemplary sensors can beused in noninvasive on-body continuous-monitoring in healthcare,fitness, remote monitoring, and other applications.

The disclosed technology includes fabrication processes to produce theexemplary temporary transfer electrochemical biosensors and chemicalsensors. For example, in some implementations, the disclosed fabricationprocesses includes a method to produce the exemplary T3 sensors that iscompatible with the non-planar features and surface irregularities thatare characteristic of the human anatomy, e.g., to provide directchemical sensing on the skin. Exemplary approaches of this method caninclude the adhesion of printable, high-resolution electrode patternsonto the epidermis using T3 substrates. The exemplary T3 chemosensorsare compatible with the skin and can be mated with and conform to thecontours of the body. In some implementations, for example, carbonfiber-reinforced tattoo inks can be employed in the exemplary T3 sensorsto provide the durability required to withstand the mechanical stressesrelevant to epidermal wear. In some examples of the disclosed T3sensors, customized artistic electrode patterns can be produced thatconceal their electrochemical functionality. The sensing paradigm of theexemplary T3 sensors can be suitable for a plethora of diverse body-wornchemosensing applications where true bionic integration is a corerequirement for developing ‘electronic skin’.

In some implementations, the disclosed electrochemical sensors andbiosensors can be printed on paper, plastic, or ceramic substrates thatare either inserted into or included on the surface of a wearable item,e.g., including, but not limited to, a wristwatch, armband, chest-strap,belt, or headband for direct epidermal contact and sensing.

In one aspect, a method of producing an epidermal biosensor includesforming an electrode pattern onto a coated surface of a paper-basedsubstrate to form an electrochemical sensor, the electrode patternincluding an electrically conductive material and an electricallyinsulative material configured in a particular design layout, andattaching an adhesive sheet on a surface of the electrochemical sensorhaving the electrode pattern, the adhesive sheet capable of adhering toskin or a wearable item, in which the electrochemical sensor, whenattached to the skin or the wearable item, is operable to detectchemical analytes within an external environment.

In another aspect, a method of producing an epidermal biosensor includesforming an electrode pattern onto a coated surface of a paper-basedsubstrate to form an electrochemical sensor, the electrode patternincluding an electrically conductive material and an electricallyinsulative material configured in a particular design layout, attachingan adhesive sheet on a surface of the electrochemical sensor having theelectrode pattern, the adhesive sheet capable of adhering to skin or awearable item and structured to include a coating layer on an externalsurface of the adhesive sheet, and removing the paper-based substratefrom the electrochemical sensor to expose the electrode pattern, inwhich the electrochemical sensor, when attached to the skin or thewearable item, is operable to detect a substance present within a fluidthat contact the electrode pattern coupled to the skin or the wearableitem.

For example, the disclosed technology can be used to construct abody-worn sensor device that is either directly attached to the skin oris included as part of a body-worn article, e.g., such as clothing, awrist band, a wrist watch, a piece of footwear, or a monitoring device.Such a body-worn sensor device can include a multi-layer materialstructure that has an electrochemical sensing material layer interactingwith a substance to be detected, an electrode layer formed on and inelectrical contact with the electrochemical sensing material layer witha printed electrode pattern to receive one or more applied electricalsignals and to output one or more electrical output signals from theelectrochemical sensing material layer indicating a reaction with thesubstance to be detected, and a base layer to provide the support to theelectrochemical sensing material layer and the electrode layer. A sensorcircuit may be integrated onto the base layer in some sensor designs, ormay be located outside the base layer, but is electrically coupled tothe electrode layer. The multi-layer material structure may be aflexible structure for attaching to skin or a body-worn object. Forskin-attached applications where such a sensor is an epidermalelectrochemical sensor device, the multi-layer material structure mayinclude a removable substrate layer such as a paper substrate and areleasing agent layer over the base layer so that device can be attachedto skin after removing the removable layer. Such an epidermalelectrochemical sensor device can use cellulose acetate, for example,over the removable paper substrate, and an insulator layer as the baselayer which can be, e.g., a silicone material such as PDMS to provide anelectrically insulating component of the overall electrode pattern ofthe sensor device.

Examples of sensor structures, materials and fabrication of the aboveand other sensor devices are provided below to illustrate variousaspects of the disclosed technology.

FIG. 31A shows a process diagram illustrating an exemplary fabricationmethod 3100 to produce epidermal electrochemical sensors of thedisclosed technology, e.g., such as the exemplary T3 electrochemicalsensors. The method 3100 includes a process 3110 to form electrodestructures 3111 on a release agent 3101 coated on a paper-basedsubstrate 3102 to form an electrochemical sensor component 3115. Forexample, the electrode structures 3111 can be patterned on the releaseagent 3101-coated paper substrate 3102, in which the electrode patternincludes an electrically conductive material and an electricallyinsulative material, and in some examples an electricallysemi-conductive material, configured in a particular design layout. Insome examples, the process 3110 can include screen printing theelectrically conductive material (e.g., electrically conductive ink) andthe electrically insulative material (e.g., electrically insulativeink), and in some examples an electrically semi-conducting ink, in thepatterned design to form the electrode structures 3111.

The method 3100 includes a process 3120 to apply an adhesive sheet 3122with a protective coating 3121 to the electrochemical sensor component3115 to form an electrochemical sensor device 3125 capable of attachingto skin (or a wearable item) for one of sensing analytes in the externalenvironment of the skin or fluids present on the skin.

The exemplary inks employed can include a wide variety of materials(e.g., such as graphite, gold, platinum, nickel, silver, silverchloride, polyethylene terephthalate (PET), polytetrafluoroethylene(PTFE), etc.), and their viscosity can be modified (e.g., using eitherbinders or solvents) as needed to yield optimal results. For example,the ink can be prepared with various chemical modifications in order toimpart selectivity, increase sensitivity, reduce response time, and/orfurther extend the stability of the amperometric, voltammetric, orpotentiometric response of the electrochemical sensor device. This caninclude the incorporation of chemical moieties into the ink suspension(e.g., catalysts, biocatalysts, enzymes, proteins, nanoparticles,reagents, mediators, binding agents, and/or cofactors), as well as thepatterning of perm-selective or ion-selective coatings/membranes to thesurface of the exemplary sensor. For example, the electricallyconductive ink can include, but is not limited to, gold, platinum,nickel, silver, and silver chloride inks. For example, the electricallyinsulative ink can include, but is not limited to, PET and PTFE inks. Insome examples, the electrode structures 3111 can also includeelectrically semi-conductive materials including semi-conductive ink,e.g., including, but not limited to, amorphous carbon, carbon black, orgraphite.

In some implementations of the method 3100, as shown in FIG. 31A, thefabricated electrochemical sensor device 3125 is prepared for epidermalmonitoring of a wearer's surrounding environment. For example, theelectrochemical sensor device 3125 could be used to detect volatileorganic compounds, explosive remnants, and pollutants present in the airsurrounding the device 3125 on the user's skin. In such cases, themethod 3100 can further include a process 3130 a to remove theprotective coating 3121 from the adhesive sheet 3122 of theelectrochemical sensor device 3125. Subsequently, the method 3100 thenincludes implementing a process 3140 a to flip the electrochemicalsensor device 3125 to be applied to skin 3141, in which the adhesivesheet 3122 is attached to the skin 3141 such that the paper basedsubstrate 3102 is positioned away from the skin 3141. The method 3100can then include implementing a process 3150 a to remove the paper basedsubstrate 3102, thereby exposing the adhered electrode structures 3111(e.g., electrode sensor pattern) to the wearer's external environmentfor remote sensing. For example, the process 3150 a can include applyingwater to the releasing agent 3101 to allow smooth release of the paperbased substrate 3102 from the electrode structures 3111.

In some implementations of the method 3100, as shown in FIG. 31A, thefabricated electrochemical sensor device 3125 is prepared for epidermalphysiological monitoring, e.g., of fluids containing biochemicalanalytes present on the skin. In such cases, the method 3100 can furtherinclude a process 3130 b to remove the paper based substrate 3102 fromthe electrochemical sensor device 3125. For example, the process 3130 bcan include applying water to the releasing agent 3101 to allow smoothrelease of the paper based substrate 3102 from the electrode structures3111. Subsequently, the method 3100 then includes implementing a process3140 b to apply the electrochemical sensor device 3125 to the skin 3141via the attachment of the adhesive sheet 3122 to the skin 3141 such thatthe external surface of the electrode structures 3111 are in contactwith the surface of the skin 3141. In some implementations, the method3100 can then include implementing a process 3150 b to remove theprotective coating 3121 from the adhesive sheet 3122 of theelectrochemical sensor device 3125.

FIG. 31B shows images 3161, 3162, and 3163 of exemplary T3electrochemical sensors showing several exemplary printed designs. Theimage 3161 illustrates a high-quality array of three-electrode artisticelectrochemical sensors possessing two varying sizes. The image 3162illustrates an array of microelectrodes that can be used forsmall-sample bioanalysis, for example. The corresponding inset in theimage 3162 exemplifies those well-defined patterns possessingmicrometer-scale resolution can be produced with the disclosedfabrication method 3100. The image 3163 shows exemplary artisticallypatterned T3 electrochemical sensors that can be employed forenvironmental sensing of the wearer's local vicinity. The exemplary T3electrochemical sensor depicted in the image 3163 indicates that theimplementation of finely-segmented and well-dispersed carbon fibers donot compromise the quality of the thick-film fabrication process. Forexample, the exemplary T3 electrochemical sensors can be fabricatedusing the described thick film process without special arrangements toaccommodate the T3 paper in the printing process.

FIG. 31C shows an image 3191 of an exemplary three-electrode T3biosensor applied to porcine skin for physiological monitoring. Theimage 3191 displays exemplary constituents of the sensor that include anAg/AgCl reference electrode (3191 a), a carbon working electrode (3191b), a carbon counter electrode (3191 c), and an insulator structure(3191 d) that circumscribes the electrodes. FIG. 31C also shows an image3192 showing a magnified view of the carbon electrodes 3191 b displayingwell-defined borders and rough morphology.

In one exemplary implementation of the method 3100, for example, a sheetof paper can be coated with a thick-film of cellulose acetate to impartrigidity for subsequent processing. After the cellulose acetate layerhas dried and solidified, a thick-film of silicone such aspolydimethylsiloxane (PDMS) can be deposited and the paper-celluloseacetate-PDMS contingent can be cured at a specific temperature tosolidify the PDMS layer. Thereafter, a thick-film of warm polyvinylalcohol (PVA) can be deposited on the surface and allowed to dry.Subsequently, the ink can be cured, e.g., at a suitable temperature. Theexemplary process is then repeated, as needed, for the number of layersrequired. Each layer can either employ an identical ink formulation asthe previous layer or a different one entirely. For example, thedescribed fabrication technique includes integration of printing andtattoo-transfer protocols and thick-film fabrication processes toproduce such advanced electrochemical sensors capable of epidermaldetection of physiologically-relevant compounds as well as agents ofenvironmental/security relevance. The exemplary technique can producebody-worn electrochemical sensors that are compliant with the skin forthe realization of non-invasive extended chemical monitoring. Nearly anyartistic tattoo design can be formed in the fabrication of theelectrochemical sensors, e.g., allowing the sensors to be concealed inrather inconspicuous tattoo artwork, without compromising the favorableresolution and performance inherent to printable sensors.

FIG. 31D shows a schematic illustration of exemplary material layers ofan exemplary epidermal electrochemical sensor device. In this example,the exemplary device includes a paper substrate and a releasing agentlayer, e.g., formed of cellulose acetate, over the paper substrate. Theexemplary device also includes an insulator layer, e.g., formed of asilicone material, such as PDMS, that is formed on the releasing agentlayer to provide an electrically insulating component of the overallelectrode pattern of the sensor device. In this example, an outercoating layer, e.g., polyvinyl alcohol (PVA), is applied on theinsulator layer. The exemplary device includes an electrode pattern,e.g., which can be screen printed on the coating layer, includingelectrically conductive and/or electrically semi-conductive inks to formthe electrodes of the sensing device in any desired pattern.

In some implementations of the method 3100, for example, carbon fiber(CF) segments can be dispersed within the tattoo ink to augment theelectrode's tensile strength and provide the electrode with aninterlinked conductive backbone while enhancing the electrochemicalbehavior, hence reflecting the inherent properties of the fiberconstituents. Inclusion of such CFs in the ink materials can counteractcracking and alleviate mechanical degradation associated with routineskin-based wear. By harnessing CF-dispersed inks for mechanicalreinforcement, the fabricated electrochemical sensors exhibitsubstantial resiliency against extreme deformation, e.g., such asrepeated pinching, bending, flexing, and twisting. The resultingwearable epidermal sensing devices of the disclosed technology thuscouple favorable substrate-skin elasticity along with highly attractiveelectrochemical performance.

FIG. 31E shows a block diagram of an exemplary embodiment of anepidermal electrochemical sensor device 3170 capable of being worn onskin or a wearable item. The electrochemical sensor device 3170 includesa substrate 3171 of an electrically insulative material, which can beconfigured as a flexible substrate. The electrochemical sensor device3170 includes a working electrode 3172 and a second electrode 3173 onthe substrate 3171, in which the working electrode 3172 and the secondelectrode 3173 are separated from one another by a spacing region 3179.The electrochemical sensor device 3170 can include an insulator layer orstructure 3176, e.g., which can provide further support for the device3170, as well as include various artistic designs like that of a tattoo.

For example, the electrode configuration of the disclosed epidermalelectrochemical sensor devices can be designed based on the type oftarget analyte to be sensed and the type of detection methodology, e.g.,amperometry, voltammetry, potentiometry, conductometry, and/orelectrochemical impedance spectroscopy, to be employed. In someexamples, the epidermal electrochemical sensor device 3170 can beconfigured to detect charged analytes, e.g., using potentiometry orconductometry. In some examples, the epidermal electrochemical sensordevice 3170 can be configured to detect self-oxidizing analytes on abare working electrode 3172, in which the device includes a thirdelectrode (not shown in FIG. 31E) positioned between the workingelectrode 3172 and second electrode 3173; and the second electrode 3173and the third electrode can serve as a counter electrode and a referenceelectrode, respectively. In some embodiments, for example, theelectrochemical sensor device 3170 includes an array of electrodes,e.g., such as an array of working electrodes, counter electrodes, and/orreference electrodes.

In other examples, as shown in the diagram of FIG. 31E, the workingelectrode 3172 includes an electrochemical sensing layer 3174 to sustaina redox reaction to produce a detectable electrical signal that can bedetected using, for example, amperometry and/or voltammetry. Theelectrochemical sensing layer 3172 provides a reaction agent (e.g., thecatalyst) that can undergo a redox reaction with a target analyte (e.g.,such as a particular molecule or substance) that produces chargecarriers sensed by the working electrode 3172. The electrochemicalsensing layer 3172 can be structured to include a catalyst and anelectroactive redox mediator. In some examples, the target analyte canbe oxidized by the catalyst, releasing electrons in the process, whichgives rise to an electrical current that can be measured between theworking electrode 3172 and second electrode 3173. For example, theelectroactive redox mediator can facilitate the transfer of electronsbetween the working electrode 3172 and the active site of the catalyst.The electrochemical sensing layer 3174 can be configured to the workingelectrode 3172 in at least one of the following configurations: (i) thecatalyst dispersed within the material of the working electrode 3172;(ii) the catalyst coated as a layer on the surface of the workingelectrode 3172; (iii) the catalyst entrapped by an electropolymerizedconducting polymer formed on the surface of the working electrode 3172;(iv) the catalyst entrapped by a selectively permeable scaffoldstructure, e.g., such as Nafion or chitosan, formed on the surface ofthe working electrode 3172; (v) the catalyst covalently bonded to thesurface of the working electrode 3172; or (vi) the catalystelectrostatically anchored to the surface of the working electrode 3172.In exemplary implementations including the electroactive redox mediator,for example, the electroactive redox mediator can be configured in theelectrochemical sensing layer 3174 along with the catalyst by the sameexemplary configuration.

As shown in the diagram of FIG. 31E, the electrochemical sensor device3170 includes an electrical sensor circuit 3177 electrically coupled tothe electrodes via electrical interconnects 3175. For example, thesensor circuit 3177 can be configured to apply excitation waveformsand/or transduce the electrical signals generated by the electrochemicalelectrodes of the electrochemical sensor device 3170 upon excitation. Insome examples, the sensor circuit 3177 can include a display or otherinterface to display the results to the wearer or other user, e.g., suchas a coach, trainer, or physician. The sensor circuit 3177 can bestructured to include, but not limited to, a potentiostat (e.g., torealize amperometric and voltammetric measurements) or a galvanostat(e.g., to realize potentiometric measurements). In some embodiments, forexample, the electrochemical sensor device 3170 can include electricallyconductive contact pads coupled to the interconnects 3175 to provide aconductive surface to electrically interface an external circuit (e.g.,such as the sensor circuit 3177) to the electrodes of theelectrochemical sensor device 3170.

The electrochemical sensor device 3170 can be applied to skin or awearable item in such a way that a sensing environment 3178 can include,for example, fluids in contact with the user's skin or clothing worn bythe user, or the external environment in which the user is in, e.g.,including air or water. The sensing environment 3178 contains the targetanalyte to come into contact with the electrodes. In some examples, ifthe sensing environment 3178 includes a fluid, e.g., such as a bodyfluid like perspiration.

Exemplary implementations of the disclosed electrochemical sensortechnology were performed, which included the described materials,procedures, and data.

The exemplary implementations described herein included the use of thefollowing materials and equipment. For example, ascorbic acid (AA), uricacid (UA), 2,4-dinitrotoluene (DNT), potassium ferricyanide (K₃Fe(CN)₆),2,4,6-trinitrotoluene (TNT), potassium phosphate monobasic (KH₂PO₄), andpotassium phosphate dibasic (K₂HPO₄) were used without furtherpurification or modification. Chopped carbon fibers (CFs), e.g.,including having 8 μm diameter, 6.4 mm length, 93% purity, wereprocessed to reduce the CF length to approximately 0.5 mm. The exemplaryreagents were prepared in a 0.1 M phosphate buffer solution (PBS, pH7.4). Ultrapure water (18.2 MΩ·cm) was used in the exemplaryimplementations, and the exemplary implementations described wereperformed at room temperature. For example, Ag/AgCl conductive inkcarbon graphite ink and insulator ink were utilized. Laser temporarytattoo paper kits were obtained from HPS Papilio (Rhome, Tex.). Forcomparison, custom-fabricated carbon screen-printed electrodes (onalumina, 2 mm working electrode diameter) were employed. Cadavericporcine skin samples were immediately refrigerated upon arrival untiltemporary transfer tattoos were applied. A CH Instruments (Austin, Tex.)model 660D electrochemical analyzer was employed, for example, for thevoltammetric, amperometric, potentiometric, and impedometricexperiments. A Keithley (Cleveland, Ohio) model 6514 system electrometerwas used to characterize trace resistance, for example. An Olympusoptical microscope with an integrated CCD camera was utilized, forexample, to investigate the surface morphology of the printed epidermalsensors in greater detail.

Exemplary sensor patterns were designed in AutoCAD (Autodesk, SanRafael, Calif.) and outsourced for fabrication on 75 μm-thick stainlesssteel stencils. For example, a separate stencil pattern was created foreach layer (e.g., Ag/AgCl, carbon, insulator). A semi-automatic screenprinter was employed for the fabrication efforts. For example, in orderto conduct electrochemical experiments, a tattoo pattern containing acircular working electrode was designed and possessed a 3 mm radius. Forexample, in order to increase the tensile strength of the printedelectrodes and mitigate the cracking observed during typical wear, 100mg of chopped CFs were dispersed in 30 mL of ink and homogenizedthoroughly.

Exemplary fabrication methods of the disclosed technology to manufactureelectrochemical biosensors and chemical sensors were employed to producethe exemplary T3 sensors for transfer onto skin (or other wearableitems) in a procedure analogous to that employed to transfer of atemporary tattoo. Exemplary techniques described involved the layeringof certain materials on a substrate (e.g., paper substrate), on top ofwhich a screen-, aerosol-, or inkjet-printed sensor pattern was defined.The substrate, e.g., containing the thick-film sensor patterns, wasreversed and applied to the skin using a damp water-infused cloth orsponge. The backing substrate of the exemplary fabricated sensor devicewas then peeled away, leaving only the functional printed sensor patternand a water-soluble synthetic polymer binder.

For example, screen printing can be employed for the formation ofthick-film electrodes intended to be used in a wide variety ofelectrochemical applications. For example, this technique employs anautomated system that guides a squeegee across a patterned stencil toextrude a specially-formulated ink in order to transfer an identicalelectrode pattern onto the substrate. This technique offers anattractive combination of moderate throughput and low cost.

Exemplary implementations were performed to investigate the printingquality of the exemplary ink materials on the T3 paper substrate. Theexemplary T3 sensors fabricated using the disclosed methods were appliedto the epidermis of various human subjects, and exemplaryimplementations were performed to evaluate the exemplary T3 sensors.

FIG. 32A shows images 3201-3204 for transferring an exemplary epidermalelectrochemical sensor device, e.g., such as the exemplary device shownin the images 3163 and 3190 of FIGS. 31B and 31C, on a user's skin. Theimage 3201 shows the exemplary electrochemical sensor device attached tothe paper substrate, in which the protective film removed and theadhesive layer exposed. The image 3202 shows the exemplaryelectrochemical sensor device with the attached paper substrate appliedto the user's skin such that the electrode patterned region is in directcontact with the epidermis. For example, the T3 paper is flipped (e.g.,electrode patterned side down) and depressed on the surface of the skin.The image 3203 shows the wetting of the releasing agent layer of theexemplary electrochemical sensor device attached to the user's skin. Forexample, the paper substrate is gently dabbed with water until itbecomes saturated. The image 3204 shows the removal of the papersubstrate, e.g., leaving the printed electrode contingent in directcontact with the epidermis. For example, the paper substrate is removedfrom the epidermis by gradually sliding it along and off the skinsurface.

FIG. 32B shows images 3205, 3206, and 3207 of several representativedesign permutations transferred onto the epidermis. The image 3205 showsan exemplary three-electrode electrochemical sensing contingent, and theimage 3206 shows the transfer of an exemplary high-resolution sensorarray onto the skin. For example, the inset image in the image 3206shows well-defined 150 μm-wide electrode features that are easilytransferred onto the epidermis, e.g., underscoring the fidelity at whichthe patterns are printed and transferred. The image 3207 displays a pairof exemplary artistically-inspired and fully functional three-electrodesensors possessing two different sizes. The disclosed T3 electrochemicalsensors can be configured in nearly any electrode design and can beimplemented without compromising the sensor functionality, e.g., such asin cases when an artistic impact is desired.

FIG. 32C shows images 3208, 3209, and 3210 that validate the structuralresiliency of the exemplary T3 sensors to extreme mechanicaldeformations, e.g., in which various strain permutations were applied tothe sensors. For example, the images 3208 and 3209 demonstratedeformation of an exemplary tattoo sensor when pinched with theforefingers or upon stretching the skin, respectively. Likewise, forexample, the image 3210 illustrates a twisting operation on theexemplary sensor. As shown in the images of FIG. 32C, the application ofthese strain permutations exhibited minimal effect on the appearance andquality of the T3 sensor. Additional exemplary implementations on theimpact of such strain permutations upon the electrochemical performanceof these printable epidermal sensors were performed.

Exemplary implementations were performed for electrochemicalcharacterization of the T3 sensing methodology aimed at comparing thedisclosed sensing paradigm with conventional screen printed electrodes(SPEs) on solid alumina substrates. For example, voltammetric signatureswere contrasted between the two systems, and a GORE-TEX fabric was usedfor the tattoo investigations in order to emulate the viscoelasticproperties of the epidermis.

FIG. 33A shows cyclic voltammogram data plots 3311, 3312, and 3313 thatwere obtained for 2.5 mM ascorbic acid (AA) at an exemplary SPE (shownin the plot 3311), at an exemplary T3 sensor on GORE-TEX (shown in theplot 3312), and at an exemplary CF-reinforced T3 sensor on GORE-TEX(shown in the plot 3313). The exemplary tattoo-based sensor embodiedfavorable electrochemical properties when compared with the conventionalSPE. As an additional benefit, for example, the incorporation of CFsinto the ink matrix enhanced the electrochemical response of the tattoosensor device substantially, leading to better-defined oxidation peaksthat emulated the response obtained at the conventional electrodecontingent. Moreover, for example, both the unreinforced T3electrochemical sensors and the CF-reinforced T3 electrochemical sensorexhibited resiliency against thirty repetitive 180° bending iterations,hence maintaining their favorable voltammetric behavior under extrememechanical strain. As can be inferred from a comparison of the insetplots in the plots 3312 and 3313, both the unreinforced andCF-reinforced T3 sensors displayed only small (e.g., less than 10%)deviations from the original current response following repetitivebending operations.

FIG. 33B shows cyclic voltammetric response plots 3321 and 3322 thatwere obtained for the detection of 2.5 mM uric acid (UA) at an exemplarySPE (shown in the plot 3321) an exemplary CF-reinforced T3 sensor onporcine skin (shown in the plot 3322). As represented from theseexemplary data, the CF-reinforced tattoo electrode exhibited improvedelectrochemical performance to that of the SPE. For example, the peakpotential and peak current assumed more desirable values when comparedwith the conventional SPE. FIG. 33B also shows an amperometric responseplot 3323 generated at the exemplary CF-reinforced T3 electrochemicalsensor for increasing UA concentration. For example, a highly linearcalibration was recorded at the skin-based electrode, corroborating itsuse not only as a viable alternative to SPEs but also as an advancedepidermal electrochemical sensor.

The disclosed tattoo sensing paradigm can also be extended to theidentification of environmental substances including hazards andpollutants present in the vicinity of the wearer, e.g., forenvironmental and security monitoring. For example, the disclosedtechnology was extended to the detection of the common explosive2,4,6-trinitrotoluene (TNT), in connection with square wave voltammetry(SWV). FIG. 33C shows SWV plots 3331 and 3332 that were obtained for theelectrochemical detection of 225 μg/mL TNT at an exemplary SPE (shown inthe plot 3331) an exemplary CF-reinforced T3 sensor on porcine skin(shown in the plot 3332). Both the plots 3331 and 3332 exemplify thewell-defined TNT response, which substantiates that the exemplary T3electrochemical sensors (even when mated with the skin) contend with theperformance offered by well-established SPE sensors fabricated on solidsupports. FIG. 33C also shows an SWV response plot 3333 of the exemplaryepidermal sensor for increasing TNT concentrations, which is shown to bewell-defined and highly linear (as shown in the inset plot).

Exemplary implementations were performed to investigate the fundamentalelectrical properties of an exemplary T3 sensor, e.g., which can beconsidered imperative in order to ascertain its utility as a viableelectrochemical device for integration with epidermal electronics. Forexample, a resistive and complex-valued impedance profile was evaluatedunder the application of mechanical deformation.

FIG. 34A shows a cyclic voltammogram plot illustrating the enhancedresponse generated by the dispersion of CF segments into the ink matrix,e.g., in which the scan rate was 10 mV/s. FIG. 34B shows a resistiveprofile plot of a normal (black squares) and carbon fiber-reinforced(red dots) 1 cm Ag/AgCl tattoo trace on porcine skin. FIG. 34C shows aplot of Nyquist complex-valued impedance curves generated by anexemplary T3 sensor before bending (black squares) and after 10 bendingoperations (red dots) on porcine skin, e.g., in which potassiumferricyanide (K₃Fe(CN)₆) was employed as the redox probe. For example,the impedance spectrogram parameters included a frequency of 0.1 Hz-10kHz, an applied potential of 0.4 V vs. Ag/AgCl, and an amplitude of 10mV_(pp).

The implementations included resistance measurements that were recordedvia the application of multimeter probes at opposite extremities of a 1cm Ag/AgCl trace on an exemplary GORE-TEX-based T3 sensor (includingboth the CF-modified and unmodified embodiments). As previously shown inthe insets of the plots 3312 and 3313, respectively in FIG. 33A, boththe exemplary unreinforced and CF-reinforced tattoo sensors exhibitedrepeatable electrochemical performance following several dozen bendingiterations. However, in this example, the unreinforced/standardelectrode trace increased in its intrinsic resistance until catastrophicfailure occurred at the 100^(th) bending iteration, e.g., represented bya completely severed trace, R=co, as shown in FIG. 34B. Conversely, forexample, in this exemplary implementation, the exemplary CF-reinforcedelectrode, although possessing slightly elevated intrinsic resistance atthe commencement of the implementation (e.g., ˜25Ω), maintained itsconductivity even following over 350 bending repetitions, and hencesubstantiating its ability to withstand highly-repetitive mechanicaldeformation and underscores its suitability for epidermal integration.

The exemplary implementations included an electrochemical impedancespectroscopy performed at the exemplary CF-reinforced T3 sensor (onGORE-TEX), e.g., to ascertain the frequency at which the compleximpedance indicates a transition from a reaction that is controlled viamass-transfer to one that is governed by kinetics. As exemplified inFIG. 34C, this transition occurred at approximately 4 Hz and 8 Hz forthis exemplary CF-reinforced T3 sensor prior to and immediatelyfollowing ten bending iterations, respectively. As shown in FIG. 34C,for example, the change in the impedance profile following stretchingwas shown to be minimal. In accordance with the Randles-Ershlerformalism, a solution resistance, R_(Ω), of ˜100Ω, charge transferresistance, R_(ct), of ˜580Ω, and a double-layer capacitance, C_(dl), of3.6 μF, can be interpolated from the plot.

The increased tolerance of the CF-dispersed electrodes against severemechanical deformation should not compromise the electroanalyticalperformance offered by the disclosed devices. This is shown in FIG. 34A.For example, an exemplary unreinforced T3 biosensor was evaluatedalongside an exemplary CF-reinforced T3 biosensor possessing a 0.04%(w/w) CF loading level in an exemplary implementation to provide thecyclic voltammetric response. As shown FIG. 34A, the electrochemicalfigures of merit differed slightly between the unreinforced (e.g.,E_(p)=0.42 V, i_(p)=15.0 μA, and k_(s)=8.8E⁻⁴ cm/s) and the 0.04%CF-reinforced sensor (e.g., E_(p)=0.32 V, i_(p)=19.0 μA, andk_(s)=1.1E⁻³ cm/s). In this exemplary implementation, the reinforcedexemplary T3 sensor exhibited more favorable electrochemical properties,as shown from the enhanced voltammetric behavior in FIG. 34A.

Exemplary implementations were performed for exemplary printed T3sensors for environmental/security monitoring applications. For example,in order to demonstrate the ability to operate in vapor-phaseenvironments, an exemplary CF-modified T3 sensor was applied towards thedetection of increasing levels of 2,4-dinitrotoluene (DNT) vapors. Assuch, the exemplary tattoo-based sensor was applied to a porcine skinsample, which was subsequently inserted in a sealed 15 mL containeralong with 100 mg of DNT salt. The system was allowed to equilibrate for30 min, after which a calibration (with respect to time) was performed.The exemplary calibration data showed a high degree of linearity alongwith rapid response time. Repeated measurements were conducted, with theexemplary resultant data demonstrating precise repeatability. Forexample, maximum 6.9% deviation in the current level at the reductionpeak (−1.05 V vs. Ag/AgCl) was observed across six independentmeasurements. For example, it is noted that vapor-phase detection istraditionally not feasible using bare SPEs due to the lack of asupporting electrolytic medium. However, the exemplary T3electrochemical sensors exhibit a noteworthy structural difference whencompared with conventional SPEs. For example, a perspiration-saturatedadhesive polymer layer can be employed as the structural backbone, whichmay behave analogous to common hydrogel layers. Thus, the exemplaryimplementations demonstrated that the exemplary tattoo-based device iswell suited to serve as a vapor-phase environmental sensor.

The subjection of conventional SPEs on rigid and flexible substrates torepeated chemical and mechanical degradation is expected to havedeleterious impact on their electrochemical behavior, thus precludingthem from epidermal integration. Advantageously, for example, thedisclosed T3 electrochemical sensors can rectify these challenges, e.g.,through the inclusion of CFs in the ink matrix as well as through theirstrongly-adhesive (and flexible) backbone. Exemplary implementationswere performed to evaluate the exemplary T3 sensors against chemicaldegradation, e.g. such as subjection to repetitive washing cycles (e.g.,t_(wash)=5 s with hand soap) to emulate hand-washing or bathing. Forexample, the washing involved generating a thorough lather with handsoap under a continuous stream of tap water for 5 s and subsequentlydrying the skin sample with a towel. FIG. 35A shows an IV data plotshowing the effect of repetitive washing cycles upon the CV waveformgenerated at the tattoo biosensor (on porcine skin) using 2.5 mM UA. Asshown in FIG. 35A, although washing did impart relatively minordegradation in the waveform, e.g., shown from the well-defined peakscorresponding to the oxidation of UA. Additionally, for example, thepeak current deviated from the baseline measurement (before washing) byno more than 15% at the conclusion of the implementations. Moreover, forexample, an increase in the oxidation current following washing mayreflect the exposure of a larger active electrode area. The peakpotential, however, remained stable throughout the course of theimplementations.

Exemplary implementations were performed to evaluate the effect ofrepetitive pinching of the tattoo patterned sensors. An exemplaryCF-reinforced T3 sensor was applied to porcine skin and repetitivelypinched for 2 s intervals, e.g., for six pinches. FIG. 35B shows an IVdata plot displaying the response of the exemplary CF-reinforced T3sensor to repetitive pinching operations employing 2.5 mM AA as a redoxprobe. The exemplary data indicate that repeated pinching of the sensorproduced minimal degradation in the electrochemical performance. Forexample, both the peak current and peak potential remained stablethroughout these pinching experiments, thereby demonstrating thecapability of high-fidelity electro-analytical operations of thedisclosed sensors under the severe demands imparted by epidermal wear.

Exemplary implementations were performed with extended durations ofroutine wear of an exemplary T3 sensor. FIG. 35C shows images 3511 and3512 captured immediately following the application of the T3 sensors onskin for an exemplary T3 sensor and an exemplary CF-reinforced T3sensor, respectively, and images 3521 and 3522 captured after 18 hoursof continuous epidermal wear of the exemplary T3 sensor and theexemplary CF-reinforced T3 sensor, respectively. A close inspection ofthe images 3521 and 3522 revealed some cracking (e.g., at theAg/AgCl-insulator and carbon-insulator interfaces) for the exemplary T3sensor without CFs, as shown in the image 3531 of FIG. 35C, andsubstantially no cracking or degradation for the exemplary CF-reinforcedT3 sensor, as shown in the image 3532 of FIG. 35C. FIG. 35C alsoincludes cyclic voltammogram data plots 3541 and 3542 for 0.5 mM UA atthe exemplary T3 and CF-reinforced T3 sensors, respectively, following12 hours of continuous wear of both sensors, e.g., with a scan rate of100 mV/s. The response data recorded at the non-CF-reinforced sensor inthe plot 3541 exhibited substantial distortion, e.g., as compared to thewell-defined anodic UA oxidation peak visible at the exemplaryCF-reinforced sensor in the plot 3541. A comparison with thevoltammograms obtained from previous implementations with UA atunperturbed electrodes also corroborated that the exemplaryCF-reinforced sensor is capable to yield high-fidelity electroanalyticalperformance over extended wear.

The disclosed technology includes techniques to the formation ofbiosensors and chemical sensors that exploit electrochemical detectionmethodologies such as amperometry, voltammetry, potentiometry, andelectrochemical impedance spectroscopy. Thus, for example, exemplarytechniques includes proper patterning of electrodes to form a completeelectrochemical system, as well as the selection of an appropriatelayering and ink formulation to facilitate the electrochemical response.The disclosed techniques can advance the field of non-invasive on-bodycontinuous-monitoring biosensors.

For example, the majority of personal blood glucose monitors rely ondisposable screen printed enzyme electrode test strips. These single-useelectrode strips are mass produced by rapid and simple thick-film screenprinting microfabrication techniques. Owing to its reliability and lowcost, the diabetic monitoring industry has leveraged this fabricationconcept for the past 30 years and has perfected the technology over thisperiod such that analytically-precise results are now achievable, evenwhen this fabrication methodology is migrated to the detection of otherphysiologically-relevant analytes such as metabolites, proteins, andDNA. The disclosed temporary transfer tattoo epidermal biosensingtechniques enable the biosensor contingent to be transferred directly tothe skin for the direct and non-invasive monitoring of the wearer'sbiochemical physiology and/or surrounding environment. For example, theexemplary biosensors can be transferred and include the ability totolerate repeated bending and stretching operations typically associatedwith on-body wear, and its extended stability on the skin. For example,this paradigm can enable continuous monitoring of the wearer'sbiochemical physiology and/or surrounding environment, in directcontrast with state-of-the art invasive “single-shot” readings such aswith blood glucose test strips for diabetics. In this manner, decreasedoverhead can be achieved, ultimately lowering the per-strip cost. Alsofor example, the biosensor pattern can be duplicated and arrayed asneeded to parallelize the sensing operation, thereby yielding asubstantially increased quantity of sensors per unit area.

In another aspect of the disclosed technology, techniques, systems, anddevices are described for fabricating and implementing tattoo-basedpotentiometric ion-selective electrochemical biosensors and chemicalsensors for epidermal and/or environmental monitoring on skin or awearable item.

For example, the disclosed tattoo-based potentiometric ion-selectiveelectrochemical biosensors and chemical sensors can includesolid-contact ion-selective electrodes (ISEs) for non-invasivepotentiometric monitoring of epidermal pH levels. The disclosedfabrication techniques of such devices include the use of temporarytransfer tattoo paper with screen printing techniques and solid-contactpolymer ISE methods. The disclosed tattoo-based potentiometric sensorsexhibit rapid and sensitive response to a wide range of pH changes withno carry-over effects. These tattoo ISE sensors are capable of enduringrepetitive mechanical deformation, which is a key requirement ofwearable and epidermal sensors. The flexible and conformal nature of thetattoo sensors enable them to be mounted on nearly any exposed skinsurface for real-time pH monitoring of the human perspiration, asillustrated from exemplary response data acquired from exemplaryimplementations during strenuous physical activity.

Potentiometric ISEs have witnessed widespread use in various researches,biomedical and industrial domains. Conventional ion-selective sensorsinclude a membrane-based ion-selective electrode and a referenceelectrode, both of which require an internal solution to ensure a stableand sensitive response. Although these sensors have been widely used invarious applications, their intrinsic design imposes inherentlimitations upon specific in vivo and ex vivo applications,particularly, for example, the internal solution complicates thefabrication process and limits their miniaturization.

The disclosed technology includes highly flexible and conformalintegrated potentiometric sensors, compatible with the non-planarity andirregularities of the human anatomy and capable of enduring prolongedmechanical strain, which can be successfully implemented in epidermalchemical monitoring, e.g., including pH measurements.

The disclosed wearable electrochemical sensing devices include aconformal geometry that is compliant with skin and can withstandrepeated mechanical stress while minimizing intrusion in the wearer'sroutine. In some implementations, the disclosed wearable electrochemicalsensing devices can be configured as textile-based sensors for in thefield monitoring of the environment and on-body monitoring, in which thetextile-based sensors conform to the wearer's anatomy while enablingunobtrusive sensing. Such wearable devices can provide detection of bothphysiological and environmental analytes. The design configurations ofthe disclosed technology enable continuous contact of detectableanalytes with the sensor surface while worn a user's body (e.g., on skinor a wearable item).

In some implementations of the tattoo-based potentiometric devices,exemplary devices include polyaniline-based solid-contact ISEs andtemporary transfer tattoo paper, and can be fabricated using hybridscreen printing techniques. FIG. 36A shows a process diagramillustrating a fabrication method 3600 to produce epidermalelectrochemical sensors with ion-selective electrodes. The method 3600includes a process 3610 to form electrode structures 3611, e.g., such ascarbon-based material electrodes, by using thick-film screen printing ona release agent layer 3601 coated over a base paper substrate 3602. Forexample, the electrode structures 3611 can be patterned on the releaseagent 3601-coated paper substrate 3602, in which the electrode patternis configured in a particular design layout. The method 3600 includes aprocess 3620 to form electrically insulative material 3626 andpoly(aniline) (PANi) 3625 to form a temporary transfer tattoosolid-contact ISE sensor 3615. The method 3600 includes a process 3630to form an adhesive sheet 3637 with a protective coating 3638 to the T3ISE electrochemical sensor component 3615 to form a T3 ISEelectrochemical sensor device 3625, which is capable of attaching toskin (or a wearable item) for one of sensing analytes in the externalenvironment of the skin or fluids present on the skin. In someimplementations of the method 3600, a process 3640 can include removingthe protective sheet 3638 from the adhesive sheet 3637 of the T3 ISEelectrochemical sensor device 3625 to enable transfer of the T3 ISEelectrochemical sensor device 3625 on a receiving surface 3649, e.g.,including skin or a wearable item An inset illustrative schematic of theapplied T3 ISE electrochemical sensor device 3625 on the receivingsurface 3649 (e.g., skin) shows the layers of materials of the exemplarydevice.

FIG. 36B shows an image of an exemplary ISE tattoo sensor including twoelectrodes, e.g., including an ISE and a reference electrode, andconnection points that can interface with a voltmeter, for example, viaelectrically conductive conduits. For example, the disclosed fabricationmethods allow development of the exemplary ISE tattoo sensors in avariety of designs, e.g. such as the ‘smiley face’ design of theexemplary ISE tattoo sensor shown in FIG. 36B. In this example, thesensor design includes one ‘eye’ of the smiley face acting as thepH-sensitive ISE while the other ‘eye’ functions as the referenceelectrode, e.g., thus concealing the complete sensor contingent in anartistic manner.

These exemplary ‘smiley face’ shaped-tattoo sensors can be readilyfabricated using tattoo base paper, electrode inks, e.g., includingcarbon and/or Ag/AgCl, and insulator inks, in which the tattoo sensorfabrication employs a distinct stencil pattern for each layer. Anadhesive sheet can later be applied to the electrode- andinsulator-printed tattoo paper for subsequent transfer on varioussubstrates. For example, poly(aniline) (PANi) exhibits pH-sensitiveconductivity, e.g., demonstrated with the reversible emeraldine salt(ES)-emeraldine base (EB) transition (acid-base reaction), and can thusbe used in the disclosed solid-state pH electrochemical sensors.Additionally, for example, PANi has minimal cytotoxicity and causesnegligible skin irritation and sensitization. Thin films of PANi can beproduced on the patterned electrodes or other structures of thedisclosed devices via electropolymerization techniques with highreproducibility, and in doing so, for example, the fabrication of theseexemplary PANi-based ISEs do not require surface pre-treatment. Thesecharacteristics along with the attractive performance make PANiwell-suited for the disclosed biocompatible, epidermal tattoo-basedpotentiometric sensors. The resulting tattoo ISE sensor devices canwithstand repeated bending and stretching operations, which are ofsubstantial relevance to wearable epidermal sensors.

Exemplary implementations of the disclosed potentiometric ion-selectiveelectrochemical sensor technology were performed, which included thedescribed materials, procedures, and data.

The exemplary implementations described herein included the use of thefollowing materials and equipment. For example, potassium phosphatemonobasic (KH₂PO₄), potassium phosphate dibasic (K₂HPO₄), hydrochloricacid (HCl), Nafion® 117 solution, aniline and citric acid were obtained.Aniline was further purified by double distillation prior to use. Carbonfibers (8 μm diameter, 6.4 cm length, 93% purity) were obtained andtheir length was reduced to ˜0.5 mm (e.g., by cutting with a sharpblade), followed by thorough cleaning in acetone. The exemplaryimplementations were conducted at room temperature, and solutions wereprepared using ultra-pure deionized water (18.2 MΩ·cm). For example,electrochemical cleaning, deposition, and potentiometric analysis wereperformed using a CH Instrument (Austin, Tex.) model 630Celectrochemical analyzer. A Mettler Toledo (Columbus, Ohio) S20SevenEasy glass-electrode digital pH meter was employed for pHmeasurements, for example. A miniaturized multimeter (Sinometer MS8216DMM) was used for on-body measurements in the exemplary implementations.

Exemplary ISE tattoo sensor devices were designed to conceal theelectrodes in a ‘smiley face’. The design included one eye functioningas the pH-sensitive ISE while the other eye functioning as the referenceelectrode. As exemplified in FIG. 36B, the two ears of the exemplary ISEtattoo sensor device were employed as connectors for attachment to adigital multimeter. For example, design of the exemplary smiley facesensor pattern was performed in AutoCAD (Autodesk, San Rafael, Calif.)and fabricated on 75 μm thick stainless steel and mesh stencils (MetalEtch Services, San Marcos, Calif.). For example, a unique stencilpattern was used for each electrode layer (e.g., including a carbonlayer, an Ag/AgCl layer, and an insulator material). For example, theconductive Ag/AgCl ink (E2414), the carbon ink (E3449), and theinsulator ink (E6165) were obtained, and an exemplary transparentdielectric ink (5036) was obtained. Carbon fiber segments were dispersedwithin the semi-conductive carbon ink matrix to increase the tensilestrength of the electrode. Printing was accomplished via an MPM SPMsemi-automatic screen printer (Speedline Technologies, Franklin, Mass.).Blank temporary transfer tattoo paper and the accompanying adhesivesubstrate were used without further derivation.

In some implementations of the method 3600, a fabrication process canfirst involve the printing of the blue insulator ink, followed by theAg/AgCl ink and the carbon ink, and finally, by another blue insulatorlayer. Following each routine, for example, the ink can be cured, e.g.,such as at conditions including 90° C. for 15 min. Subsequently, forexample, a 30 wt % KCl-doped transparent insulator can be screen printedonly on the surface of the reference electrode and then cured, e.g.,such as at conditions including 90° C. for 6 min. Finally, for example,a total of 6 μL of the 5% Nafion solution can be drop-casted on theAg/AgCl reference electrode and left to dry, e.g., for an overnightduration. This exemplary fabrication process was implemented to producean exemplary ISE tattoo sensor device used in subsequent implementationsdescribed.

In some implementations of the fabrication process, for example, priorto the electropolymerization of the aniline material, the workingelectrode was electrochemically cleaned by five cyclic voltammetricscans in 0.5 M HCl over the potential range of −0.3 V to 1.1 V (e.g., anexternal Ag/AgCl reference electrode and an external Pt wire auxiliaryelectrode were used in this processing step). Surface modification withPANi was performed in a 0.1 M aniline/1 M HCl solution by cyclicvoltammetry from −0.2 V to 1.0 V (vs. Ag/AgCl) at 0.1 V/s. In suchexamples, electropolymerization was first performed for 12 cycles, thena fresh solution was dispensed on the surface, followed by additional 13cycles. A total of 25 cycles were thus executed for the completepolymerization of the working electrode surface. During the exemplarycleaning and polymerization steps, in this example, the screen printedAg/AgCl electrode was protected from the electrolyte solution to avoidits damage by highly acidic solutions and aniline. After air-drying theexemplary PANi film, the adhesive sheet was applied to the tattoo. Forproper contact between the two electrodes and analyte solution, thisadhesive sheet was excised to remove a rectangular-shaped region aroundthe two electrodes (e.g., the two eyes). The as-prepared ISE tattooswere then ready for transfer and evaluation.

Exemplary implementations of the fabricated ISE tattoo sensor deviceswere examined in vitro by applying them onto hard plastic substratesprior to on-body epidermal studies. In some exemplary implementations,the tattoo ISE sensors were analyzed within the pH range of human sweat(e.g., pH 3-7, with a mean around pH 5) using standard Mcllvaine'sbuffers. For example, since human perspiration can exhibit continuousfluctuations of pH, a practical pH sensor must encompass a rapid andnear-instantaneous response to pH modulations over this range.

FIG. 37A shows a data plot of the potential-time response of anexemplary ISE tattoo sensor for decreasing pH levels and an inset plotof electrical potential versus pH, e.g., using the standard McIlvaine'sbuffers. The data plot of FIG. 37A displays a characteristicpotential-time recording at the exemplary tattoo-based potentiometricsensor for decreasing pH levels between 7 to 3 (in one-unit decrements).This real-time recording illustrates that the disclosed ISE tattoosensors exhibit a nearly instantaneous response to varying pH solutions,e.g., yielding 80% of their steady-state signal within the first 10 secwhile a completely stabilized signal was observed within 25 sec. Theresulting calibration plot (shown in the inset of FIG. 37A) displays asub-Nernstian behavior, e.g., with a mean slope (s_(x)) of 50.1 mV/pHand a relative standard deviation (RSD) of 3.72% (n=4). The pHsensitivity (slope) and conductivity of PANi depend on orientation ofthe crystalline and amorphous phases of PANi. The observed sub-Nernstianresponse of the PANi tattoo sensors can be attributed to inferiororientation of these phases. As discussed later, mild mechanicaldeformations to the ISE tattoo sensor devices caused reorientation ofthe conducting and amorphous phases and improved the pH-sensitivity to anear-Nernst response. Batch-to-batch variations between the tattoos alsoexhibited a low RSD of 4.63 (n=4), hence indicating the capability ofthe described fabrication techniques to produce reproducible devices.

It has been observed that the pH of human perspiration fluctuatesaccording to the respiration rate. As such, the exemplary ISE tattoosensor devices must also exhibit minimal carry-over in order to monitorsuch dynamically-fluctuating pH environments. To investigate thisparameter, the ISE tattoos were subjected to operation in varying pHsolutions and consecutive measurements recorded without reconditioningor rinsing of the tattoo surface.

FIG. 37B shows a data plot of the potential-time response of theexemplary ISE tattoo sensors, which demonstrates the reproducibility ofthe sensors in response to large pH fluctuations. The data plot of FIG.37B demonstrates the dynamic response of the exemplary tattoo ISE sensordevices to alternate and multiple exposures to solutions of differentpH. The exemplary device responded rapidly and favorably to thesedynamic pH changes, regaining rapidly the same potentiometric signal fora given solution pH during this continuous operation. The negligiblecarry-over of the exemplary tattoo-ISE response reflects the fact thatthe emeraldine salt (ES)-emeraldine base (EB) transition of PANi is fastand reversible. Thus, the tattoo sensors have the capability to performeffectively under continuously-varying pH milieu, viz., in situ pHmeasurement of human perspiration with low carry-over.

A distinctive feature of wearable sensors is their ability to endureprolonged mechanical strain, which is a key requirement of wearable andepidermal sensors. This is especially true in the sports, athletics,fitness and military domains. Exemplary implementations of thefabricated ISE tattoo sensor devices were performed to examine theinfluence of relevant mechanical stress upon the sensor performanceprior to their integration with the epidermis.

For example, the influence of mechanical strain permutations, includingrepeated bending and stretching, upon the potentiometric response wereexamined. In some examples, the exemplary ISE tattoo sensor devices weresubjected to a total of 50 bending and 40 stretching applications. Forthese exemplary implementations, the tattoos were transferred ontoGORE-TEX as its viscoelastic behaviour mimics that of skin. In thebending implementations, the tattoo was bent to 180° and maintained atthat position for 5 sec prior to release. The response of the tattoo wasmeasured subsequent to 10 bending iterations from pH 7 to 3. The effectof stretching upon the electrochemical performance of the tattoos wasanalyzed. In the stretching implementations, the exemplary ISE tattoosensor devices were stretched an additional 10% in lateral extent andmaintained at that position for 5 sec followed by release andinvestigation of the response. In cases of stretching deformation, theresponse was measured at an interval of 5 consecutive stretchingoperations.

FIGS. 38A and 38B show a data plot and images representing the influenceof repeated mechanical strain (e.g., bending) upon the response of anexemplary tattoo ISE device. The data plot in FIG. 38A shows thepH-responsive behavior of the exemplary ISE tattoo sensor over the 3-7pH range prior to stretching (black waveform 3801 and black squares inthe inset plot) and following the 50^(th) bending on GORE-TEX (redwaveform 3802 and red dots in the inset plot), e.g., one unit pHdecrement per addition. The images of FIG. 38B show the exemplary tattooISE sensor device applied to the cubital fossa at 0° bending, 90°bending, and after the 50^(th) bending.

FIGS. 39A and 39B show a data plot and images representing the influenceof repeated mechanical strain (e.g., stretching) upon the response of anexemplary tattoo ISE device. The data plot in FIG. 39A shows thepH-responsive behavior of the exemplary ISE tattoo sensor over the 3-7pH range prior to stretching (black waveform 3901 and black squares inthe inset plot) and following the 40^(th) bending on GORE-TEX (redwaveform 3902 and red dots in the inset plot), e.g., one unit pHdecrement per addition. The images of FIG. 39B show the exemplary tattooISE sensor device applied to the forearm at normal, after the 1^(st)stretch, and after the 10^(th) stretch.

It is noted, for example, the deformation created during these exemplaryimplementations of the exemplary tattoo ISE sensor device utilizedyielded a beneficial effect upon the response of the sensor.Specifically, for example, in the absence of applied deformation, theexemplary tattoo ISE sensor device yielded a sub-Nernstian response(e.g., 52.8 mV/pH), e.g., as observed with plastic substrates. Theresponse of the exemplary tattoo ISE sensor device improved to 59.6mV/pH within the first 10 bending iterations (as shown in the data plotof FIG. 38A). Thereafter, the response stabilized to yield a final slopeof 57.5 mV/pH following the 50^(th) bending iteration. The RSD for theentire exemplary implementation was 5.71%. A similar trend was observedfor the stretching implementations, e.g., where an initial slope of 53.0mV/pH increased to 58.2 mV/pH following the 10^(th) stretching iterationand finally stabilized at 57.54 mV/pH after the 40^(th) stretch (asshown in the data plot of FIG. 39A). A 4.72% RSD was obtained in thisexemplary case. The sensitivity enhancement observed may be attributedto uncoiling and reorientation of the crystal and amorphous phases ofPANi and the subsequent improvement in its conductivity owing tomechanical deformation.

Visual analysis of the exemplary tattoo ISE sensors under bending andstretching were performed on the human skin. For the bending studies,for example, the exemplary tattoo ISE sensor device was applied to thecubital fossa and the arm was bent completely until the fingers touchedthe scapula acromion, thus simulating the extreme deformation expectedunder heavy epidermal wear (as shown in the images of FIG. 38B). In thestretching scenario, for example, the exemplary tattoo ISE sensor devicewas applied to the forearm and then stretched repeatedly to the maximumextent (as shown in the images of FIG. 39B). These images reveal thatthe potentiometric sensors are quite resilient and that their structuralintegrity does not easily degrade. Accordingly, the exemplary tattoo ISEsensor device are well-suited for applications involving continuousmotion of the substrate, e.g., as normally experienced by the humanbody.

There are growing demands for ion-selective sensors in the medical,sports, athletics, and fitness fields where point-of-care devices forthe monitoring of physiological conditions are required. Electrolytes(e.g., such as Na, Cl, K, and/or Mg) and pH levels of perspiration canreadily yield information regarding the metabolic state of an individualas well as their respiration dynamics during a fitness routine. Thus,continuous pH analysis of human perspiration is of great importance inthe areas of clinical diagnostics and sports medicine.

Exemplary implementations of the exemplary tattoo ISE sensor deviceswere performed to demonstrate such continuous, real-time physiologicalmonitoring under various conditions. Exemplary tattoo ISE sensor deviceswere applied to different locations throughout the body (e.g., the neck,wrist and lower back) of an active, consenting volunteer (sex: male;age: 27; weight: 70 kg; height: 186 cm) while the multimeter readoutunit was attached to the body using a commercially-available arm-band.For example, the multimeter leads were attached firmly to the connectionpoints (e.g., the designed ‘ears’) of the exemplary tattoo ISE sensorsusing transparent tape. The elasticity of the tattoo substrate allowsthe exemplary potentiometric sensor to attach firmly to these differentbody locations. In these exemplary implementations, for example, theneck, wrist, and lower back areas were selected in order to varyoperational conditions experienced by these sensors, as mechanicalstress and local pH are expected to vary among these locations.

For example, the ISE application process and the subsequent mating ofthe sensor with the readout instrument required less than 5 min and wasreadily performed by the subject. During the experiment the subject useda stationary cycle in a gymnasium for a total of 40 min followed by a 10min gradual cool down. The subject ingested no fluid (dehydrated state)during the entire exercise. Heart rate and cadence were maintainedaround 165 and 130 RPM, respectively. pH sensing of the subject'sperspiration was performed by the exemplary ISE tattoo sensor devicesand the data were collected at regular intervals using the miniaturizedmultimeter. To confirm that the tattoos yielded accurate readings, theregional pH was verified using a conventional pH meter and glasselectrode.

FIG. 40A shows a data plot of the real-time voltage-time response of anexemplary ISE tattoo sensor applied to a subject's neck to detect pHchanges (shown as waveform 4011), e.g., as compared to that of aconventional pH meter (with response shown as waveform 4012).

FIG. 40B shows a data plot of the real-time voltage-time response of anexemplary ISE tattoo sensor applied to a subject's wrist to detect pHchanges (shown as waveform 4021), e.g., as compared to that of aconventional pH meter (with response shown as waveform 4022).

FIG. 40C shows a data plot of the real-time voltage-time response of anexemplary ISE tattoo sensor applied to a subject's lower back to detectpH changes (shown as waveform 4031), e.g., as compared to that of aconventional pH meter (with response shown as waveform 4032).

FIG. 40D shows an image showing the exemplary device used in theseexemplary implementations (e.g., the tattoo ISE sensor interfaced with adigital multimeter) attached to the subject's wrist for the epidermalmeasurements of pH in human perspiration.

Although an athlete initially perspires at a low rate, this is soonfollowed by heavier perspiration as physical activity continues. Thus,an important requirement for such tattoo sensors is their ability toyield precise readings during a wide range of sweat flow rates. It wasobserved that during the first 10 min of exercise, the exemplary ISEtattoo sensor devices provided no response as the amount of perspirationgenerated was not sufficient to record a consistent open circuitpotential. This was also true for the pH glass electrode. However, atthe 10 min mark, sweat excretion became sufficient for the tattoosensors to yield a stable reading. Initially, the pH measured at thethree positions (e.g., the neck, wrist, and lower back) by the tattooISE sensors were almost the same (e.g., ˜pH 5.3), as shown in FIG.40A-40C.

In this exemplary implementation, the real-time sweat pH data obtainedfrom the exemplary ISE tattoo sensor devices can be explained based onvarying sweat rate at the respective body parts. The exemplary subjectperspired most profusely in the vicinity of the neck, followed by thelower back and the wrists. As the sweat excretion rate increases, therelative concentration of lactate and pyruvate decreases due todilution, and the pH concomitantly increases. The data plots of FIGS.40A-40C illustrate that the exemplary ISE tattoo sensor devicesperformed favorably with a mean slope of ˜54 mV/pH, and theirpotentiometric response at the different body locations followed closelythe pH values recorded with the glass electrode.

During the entire course of the exemplary implementation, it wasobserved that the exemplary ISE tattoo sensor devices performed wellduring both moderate and profuse perspiration. However, owing to thecombination of excessive sweating and the highly curvilinear morphologyof the skin on the neck, the neck-based tattoo ISE sensor functionedreliably for about 30 min. It is also noted that the exemplary ISEtattoo sensor devices functioned satisfactorily even when minor crackswere observed (as long as connection to the multimeter was maintained).This can be attributed to the fact that the potentiometric response isindependent of electrode area, e.g., which is in contrast toarea-dependent voltammetric and amperometric type measurements.

In another aspect of the disclosed technology, techniques, systems, anddevices are described for fabricating and implementing temporarytransfer tattoo-based electrochemical biosensors for non-invasivemonitoring of lactate in perspiration, e.g., in which the sensors areapplied to skin or a wearable item.

Lactate is a key stress biomarker and has garnered substantial interestin the athletics field. Muscular fatigue is a major hindrance in anathlete's performance and thus extensive efforts are taken to improveone's stamina. This is especially true in intensive and endurance-basedsports such as the triathlon, cycling, boxing etc. Lactate is widelyrecognized as an important biomarker of muscular exertion and fatigueand has been extensively utilized by coaches, exercise physiologists,and sports physicians to monitor an athlete's performance. When anindividual engages in intense physical activity, the body makes thetransition from aerobic to anaerobic respiration in order to satisfy theenergy demands of the musculature. During this phase, the body consumesstored glycogen from the muscles to generate energy; an unwanted effectof this process is the production of lactate, which is associated with aburning sensation in the muscles. This process is known as “glycolysis”or “lactate acidosis”. During glycolysis, the lactate levels inperspiration increases, as well. There exists a correlation betweenperspiration lactate and blood lactate, and therefore, perspiration canbe used as the sample for the analysis of muscular exertion and fatiguein persons without the need for finger sticks or venipuncture.

The disclosed technology includes fabrication methods to producetemporary transfer tattoo electrochemical sensing devices fornon-invasive enzymatic detection and quantification of lactate in humanperspiration. Implementation of the disclosed T3 sensors can providecrucial insight into an athlete's metabolic response to controlledphysical activity by offering critical insight into the temporaldynamics of lactate concentration in the perspiration. Currently, thegold-standard for lactate monitoring in the fitness, athletics, andsports domains is the use of blood lactate sensors, which areenzyme-functionalized electrochemical strips requiring finger stickblood samples, akin to blood glucose readings. The present blood lactatesensors show high sensitivity and selectivity towards lactate and aresuccessful in detecting it within 5-15 seconds. However, a majordrawback inherent to these sensors is their invasiveness and samplecollection methodology. Furthermore, to obtain a detailed lactateprofile, the blood is usually collected at a coarse interval of fewminutes while the athlete engages in rigorous training, which invariablyhinders performance. A non-invasive lactate sensor offering highertemporal resolution is thus highly desired.

The disclosed techniques for real-time non-invasive lactate sensing inhuman perspiration use the described printed electrochemical temporarytransfer tattoo biosensors. These exemplary enzymatic T3 electrochemicalbiosensors are capable of adhering to the epidermis and demonstrateresiliency against continuous mechanical deformations common toepidermal wear. The biosensors can be implemented for real-time, on-bodyanalysis for the detection and quantification of lactate inperspiration, e.g., during fitness and activity to provide usefulinsight into a user's health and athletic performance, in which thesensed data can be used for enhancing the athletic performance of theuser.

In some implementations, an exemplary enzymatic T3 electrochemicalsensor device includes a working electrode, a counter electrode, and areference electrode configured on a flexible electrically insulativematerial structured to adhere to the skin (or a wearable item) of auser. FIG. 41A shows a schematic illustration of an exemplary enzymaticT3 electrochemical sensor device including three electrodes (e.g., theworking electrode, counter electrode, and reference electrode)configured in the design ‘tattoo’ design “NE” for electrochemicaldetection of L-Lactic acid.

The working electrode of the exemplary T3 sensor can be functionalizedwith monolayers, ligands, enzyme catalysts and/or electroactive redoxmediators, among other molecules or substances to enhance thedetectability of the target enzyme. For example, in the exemplary T3sensor of FIG. 41A, the working electrode is functionalized withtetrathiafulvalene (TTF), an electroactive redox mediator, andmulti-walled carbon nanotubes in order to tether the active site oflactate oxidase, e.g., an enzyme catalyst, to form an electrochemicaltransducer layer on the electrode surface of the working electrode. Alsoin this example, a layer of chitosan is deposited onto theenzyme-electrode to impede the efflux of the biocatalytic backbone fromthe electrode to the aqueous environment. FIG. 41B shows a schematicillustrating an exemplary modified working electrode including thetransducer layer coated by biocompatible polymer (e.g., chitosan). Underthis scheme, lactate diffuses through the chitosan membrane and isoxidized by LOx to pyruvate, releasing two electrons in the process,which give rise to an electrical current that can be measured betweenthe working and counter electrodes.

For example, during bouts of physical exertion, the human body performscomplex motions that cause the skin to undergo extreme mechanicaldeformations. Hence, wearable devices must survive such harsh conditionswithout compromising their performance. The disclosed T3 electrochemicalbiosensors possess the capability to withstand repeated iterations ofmechanical deformation. Additionally, the disclosed T3 electrochemicalbiosensors possess the specificity to detect the target analyte oranalytes desired, e.g., from human perspiration. FIG. 41C shows an imageof the exemplary enzymatic T3 sensor device transferred on human skin.

Exemplary implementations were performed using the disclosed enzymaticT3 electrochemical sensor devices applied to skin of human subjects forin-situ sweat lactate profile recordings and analysis. A comparison ofthe profiles with previously reported data substantiate that the lactateepidermal biosensor platform performs desirably under conditions typicalof use, e.g., demonstrating their utility as a non-invasive technique toassess lactate levels in order to assess and affect physicalperformance.

The exemplary implementations of the disclosed enzymatic T3electrochemical sensor technology included the described materials,procedures, and data.

The exemplary implementations described herein included the use of thefollowing materials and equipment. For example, tetrathiafulvalene(TTF), glutaraldehyde solution (8%), chitosan, acetic acid, bovine serumalbumin (BSA), L-lactic acid, sodium phosphate monobasic (NaH₂PO₄),sodium phosphate dibasic (Na₂HPO₄), D(+)-glucose, L(+)-ascorbic acid,uric acid and creatinine were obtained and utilized in the exemplaryimplementations. Additionally, L-Lactate oxidase (LOx) andcarboxy-functionalized multi-walled carbon nanotubes (MWNTs) were alsoacquired and used. Exemplary reagents were used without furtherpurification. Carbon fibers (CFs) (e.g., 8 μm diameter, 6.4 mm length,93% purity) were obtained and further processing was performed to reducetheir length to approximately 2 mm, for example, and the CFs werecleaned with acetone. Electrochemical characterization was performed atroom temperature leveraging a CH Instruments (Austin, Tex.) model 1232Aelectrochemical analyzer.

Exemplary enzymatic T3 electrochemical sensor devices were designed inthe shape of “NE” (e.g., acronym for “NanoEngineering”). An exemplaryfabrication process, e.g., similar to that described in FIG. 36A, wasemployed. Briefly, for example, the fabrication of the exemplary“NE”-designed enzymatic T3 sensors included dispersing chopped carbonfibers within both semi-conductive carbon (E3449) and conductive silverink (E2414), e.g., to a concentration of 1.5% and 1.2%, respectively, toincrease the tensile strength of the electrodes. Corresponding stencilpatterns were designed and used for printing each layer on tattoo basepaper, following the sequence of carbon, silver and insulator, using anMPM-SPM semi-automatic screen printer (Speedline Technologies, Franklin,Mass.). As shown in FIG. 41A, the ‘E’ portion of the exemplary deviceincludes a reference electrode (e.g., fabricated from the exemplarysilver ink), and a counter electrode and a working electrodes (e.g.,fabricated from the exemplary carbon ink). A transparent insulator wasscreen printed on top to confine the electrode areas. Following everyscreen printing step, the printed tattoo paper was cured at 90° C. for15 min in a convection oven.

Upon the fabrication of the exemplary NE″-designed enzymatic T3 sensordevice, the working electrode was further functionalized. For theseexemplary implementations, MWNTs were first suspended in ethanol (e.g.,5 mg/mL), and sonicated for several hours until uniform suspension wasachieved. The suspension was then mixed with 0.1 M TTF ethanol/acetone(e.g., 9:1 (v/v)) solution in a volume ratio of 2:1 and sonicated for 1h. For example, 3 μL of MWNTs/TTF suspension was subsequently cast ontothe open area of the working electrode. After the electrode completelydried, 3 μL of LOx solution (e.g., 40 mg mL⁻¹ with 10 mg mL⁻¹ BSA) wascast on the electrode and dried under ambient condition, and latercovered with 2 μL of 1 wt % chitosan solution. The electrodes were thencross-linked with glutaraldehyde vapor overnight at 4° C.

The transfer process of the exemplary NE″-designed enzymatic T3 sensordevice on a host surface (e.g., including human skin) was similar to thepreviously described process, for example, as in FIG. 36A, with minormodifications. For example, in these exemplary implementations, a voidwas maintained around electrode areas to facilitate the flow ofperspiration among electrodes for the on-body tests. For the exemplaryin vitro implementations, the exemplary NE″-designed enzymatic T3 sensorwere applied to the host surface such that the bio-functionalized sidefaced upwards, while during the exemplary on-body implementations, thebio-functionalized side faced downwards (in direct contact with humanskin).

The exemplary implementations included evaluating the electrochemicalperformances of the exemplary lactate T3 sensor in vitro by transferringit onto a rigid plastic substrate and/or onto a flexible GORE-TEXtextile for mechanical integrity studies. These analyses were performedusing 0.1 M phosphate buffer, pH 7.0. For example, the operationpotential for the exemplary lactate T3 sensor was determined in vitro byapplying linear sweep voltammetry with a scan rate of 1 mV/s from −0.2to 0.2V using 8 mM L-lactate. The amperometric responses were recordedat a constant potential of 0.05 V for 60 s after 1 min incubation. Inthe exemplary stability implementations, for example, amperometricresponse to 8 mM L-lactate was conducted every 30 min for an 8 h-period.In between the exemplary implementations, the tattoo was kept at roomtemperature. For exemplary interference assessments, potentialinterferents with average concentration existing in human sweat wereexamined.

The exemplary implementations included evaluating the electrochemicalperformances of the exemplary lactate T3 sensor in on-body epidermalL-lactate sensing applications. For example, the exemplaryimplementations included healthy subjects asked to wear a tattoo lactatesensor on their deltoid in order to assess real-time lactate generation.The sensor was connected to the CHI analyzer using fine stainless steelwires, and the real-time lactate profile was recorded using amperometry(time interval: 5 s, potential: +0.05V). Subjects were asked to mount astationary cycle, begin cycling at a steady, slow cadence for 10 min.Following this ‘warm-up’ period, subjects were instructed to cycle withan increasing resistance every 3 min until maximum he/she can reach.This process ensured that the anaerobic respiration threshold wasattained, hence augmenting the excretion of lactate in the perspiration.Later subjects were asked to gradually reduce their cadence during a 10min ‘cool-down’ period. The exemplary subjects ingested no fluid duringthe duration of the fitness routine. During the workout, blood lactateconcentrations were measured using a commercial lactate sensor. Thecorrelation between sweat and blood lactate concentration were analyzed.

As shown in FIGS. 41A and 41B, the exemplary lactate T3 electrochemicalsensors included a functionalized working electrode forming aMWNTs/TTF/LOx/Chitosan matrix. For example, during high body activity,e.g., caused during sports, fat nutrition, stress, infections and/ororgan malfunction, the usual aerobic metabolism is incapable ofsatiating the energy needs of the human body. In such times, theanaerobic process (glycolysis or lactate acidosis) is initiated whereinthe stored glycogen is consumed to produce energy and lactate by musclecells. The sweat lactate concentration is a function of glycolysis andsweat rate and changes continuously with time. Moreover, the lactateconcentration of the human sweat depends on a person's metabolism andcan vary between 3 mM to 50 mM. However, in most cases the sweat lactateconcentration fluctuates within 3-25 mM. Thus a wide linear detectionrange coupled with fast response time is mandatory for an ideal sweatlactate sensor.

For example, in the case of lactate, typical lactate sensors are oftenbased on two types of enzymes, lactate dehydrogenase and lactateoxidase. However, one must recognize that for LDH, NAD⁺ must be employedas the cofactor, which represents a noteworthy challenge given that thismolecule must be immobilized on the electrode to prevent it fromleaching into the solution while being able to diffuse, with relativeease, to the enzyme's active site. The detection of lactate from LOx isusually at a high potential (>+0.65V). At such high potentials, otherelectroactive metabolites become active and lead to false data.

In the disclosed technology, for example, mediators, e.g., such as TTF,are used in the exemplary T3 electrochemical biosensors to achieveelectrocatalytic conversion of L-lactate by LOx at lower potential,e.g., thus avoiding interference by other electroactive species. Tofurther improve the efficiency of the exemplary lactate T3 sensors,MWNTs are dispersed together with mediator to serve as the electrontransducer on the working electrode. Furthermore, given the aim ofepidermal usage of the tattoo sensor, the tattoo was coated with abiocompatible chitosan layer that functions as physical barrier andlimits the efflux of the catalytic backbone from the tattoo and onto theunderlying skin. This exemplary functionalization scheme of the workingelectrode of the exemplary lactate T3 electrochemical sensor is shown inFIG. 41B.

Exemplary implementations of the exemplary lactate T3 electrochemicalsensor devices were performed in in vitro applications. For example,linear sweep voltammetry was applied first in the presence and absenceof L-lactate in buffer. In this example, the exemplary lactate T3 sensorshowed a peak value around +0.05 V, e.g., indicating that theMWNTs/TTF/LOx/Chit exhibits selective catalytic ability towards theoxidation of L-lactate. The potential of +0.05 V was applied for all thefollowing amperometric detections. The exemplary lactate T3 sensors werethen implemented, for example, to identify the detection range and theresponse time. In this exemplary implementation, the LOx functionalizedT3 devices were exposed to varying concentrations of lactate preparedusing, for example, 0.1M sodium phosphate buffer (e.g., pH 7.0).

FIG. 42A shows an amperometric data plot of the responses for differentconcentrations of L-Lactate using the exemplary lactate T3electrochemical sensor device, e.g., with 1 mM increment,E_(applied)=+0.05V. FIG. 42B shows the exemplary correspondingcalibration plot of L-Lactate. As shown in FIG. 42A, the exemplarylactate T3 electrochemical sensors exhibited linear detection from 1 mMto 20 mM beyond which the signal gradually saturates with 644.2 nA/mMsensitivity and correlation (e.g., current(μA)=0.644[L-lactate](mM)+0.689) between current and lactateconcentration. For example, the large detection range and highsensitivity may be attributed to the high surface area provided by theMWNTs which can augment the enzyme loading capacity of the tattoos andthe fast electron transfer between enzyme and transducer.

For example, a person can apply an exemplary lactate T3 sensor of thedisclosed technology to the epidermis and continuously monitor thehealth status. For such uses, the lactate T3 sensors can provide stablereproducible signals at room temperature during long periods ofoperations. The stability of the lactate T3 sensors may depend on thestability of the enzyme. To evaluate this, a time dependent analysis ofan exemplary lactate T3 sensor device was performed in which theresponse of the tattoo was recorded. FIG. 43A shows a data plot showingthe stability of an exemplary lactate T3 electrochemical sensor. Theinset plot shows the corresponding current-time data of the amperometricresponses. The data plot of FIG. 43A shows that the exemplary sensorprovided reproducible results (e.g., RSD=3.6%), which underscores itsapplicability for long term epidermal use.

In addition, for example, the human sweat includes several metabolitesand electrolytes. Out of these, creatinine, ascorbic acid, glucose anduric acid can affect the response of the exemplary enzymatic T3 sensors.An exemplary lactate T3 sensor device was implemented in presence ofthese exemplary interferents at physiological concentrations. FIG. 43Bshows a data plot of an interference implementation with (a) 4 mML-lactate, (b) 84 μM creatinine, (c) 10 μM ascorbic acid, (d) 0.17 mMglucose and (e) 59 μM uric acid. As shown in the data plot of FIG. 43B,the exemplary interferents exhibited minimal effect on the response dueto lactate with signal deviation not more than 6% for each of theinterferents.

For example, the human epidermis regularly experiences deformations dueto bodily movements. Such epidermal deformations are a major cause ofconcern for wearable devices in which the devices undergo disfigurationssimilar to the skin. This can be true for epidermal electronics sincethese go directly on the human skin. As the human body moves, the skincan undergo bending, stretching and twisting stress. Mechanical straincan lead to increased surface area of printed wearable devices. Theamperometric response is a function of the electrode area. Varyingelectrode area affects the sensor signal and can lead to undesiredresults.

Exemplary implementations were performed to evaluate the mechanicalresiliency of the exemplary lactate T3 electrochemical sensor devices.The robustness of the exemplary T3 sensor was implemented by applying itto GORE-TEX and bending it for 120 times by 90° while the sensorresponse was recorded after every 20 bending iterations. This wasfollowed by stretching the same tattoo by 10% for a total of 80 timeswith data recorded every 10 stretching iterations. Eachbending/stretching cycle included bending/stretching for 5 s followed byrelaxation of another 5 s. FIGS. 44A-44D show data plots of theelectrochemical responses of an exemplary lactate T3 sensor transferredon a flexible GORE-TEX textile undergoing repeated bending (FIG. 44A)and stretching (FIG. 44C), with their normalized current plots (FIGS.44B and 44D, respectively). As demonstrated by the data in the dataplots of FIGS. 44A-44D, the response of the exemplary lactate T3 sensorin each mechanical stress implementation remained substantially stablewith an exemplary R.S.D. of 1.24% and 1.50% during bending andstretching, respectively. For example, the minimal deviation of theexemplary sensor response even after subjecting it to large number ofstress cycles may be attributed to two reasons, e.g., (i) the carbonfiber dispersed in the carbon and Ag/AgCl inks and the MWNTs drop castedon the printed T3 sensor may enhance the resiliency of the tattoostowards mechanical deformations while providing electrical connectivityand (ii) the transparent insulator covering majority of the exemplary T3sensor surface area may further help in avoiding crack developmentswithin the device. Therefore, the disclosed enzymatic T3 sensors are acapable of performing desirably under various strains and thus serve asa compelling epidermal sensing platform.

FIG. 45 shows images of exemplary lactate T3 sensors on human skin onthe neck under mechanical strain including stretching (top row), bending(middle row), and twisting (bottom row) endured by a bare ‘NE’ tattooduring and subsequent to 100 stretching, bending, and twistingiterations (shown in the right column of images). The exemplary imagesdemonstrate that the T3 sensors are quite resilient to such flexions.

The disclosed epidermal lactate T3 sensors can be implemented forreal-time, online monitoring of the lactate concentration during humanexercise. Examples of screen-printed templorary transfer tattoo lactatesensors were modified with LOx for the oxidation of sweat lactate andMWNTs/TTF to enhance transduction, which were used in various in vitroand on-body implementations. In the exemplary in vitro implementations,the exemplary enzymatic T3 sensors exhibited a wide linear range up to20 mM with a high sensitivity of 644.2 nA/mM. The exemplary enzymatic T3sensors also showed specific selectivity toward lactate, e.g.,demonstrated in exemplary implementations that included adding severalinterfering metabolites common in sweat. In addition, for example, theelectrochemical performance of the exemplary enzymatic T3 sensors wasshown to be consistent with bending and stretching the tattoostransferred on GORE-TEX and human skin located at the neck. Exemplaryresults showed that the exemplary enzymatic T3 sensors can providereal-time sweat lactate concentration patterns.

In another aspect of the disclosed technology, the disclosed epidermalelectrochemical sensors can be included with epidermal biofuel cells toform an on-body, wearable complete self-powering monitoring system.

Examples of the epidermal printed biofuel cells including methods,systems, and devices are described in the PCT Patent Applicationdocument, entitled “PRINTED BIOFUEL CELLS”, filed Nov. 30, 2012, whichis incorporated by reference in its entirety as part of the disclosurein this patent document.

Exemplary implementations of the printed biofuel cells and methods tofabricate them are described in this patent document.

The disclosed technology includes wearable epidermal biofuel celldevices to provide continuous power generation while worn on a human orother user. In some implementations, the exemplary wearable biofuel celldevice can be applied to the wearer's epidermis as a temporary-transfertattoo and is able to scavenge an ample supply of the biofuel L-lacticacid found in the wearer's perspiration in order to generate power. Inthis exemplary device, the electrodes of the wearable epidermal biofuelcell can be functionalized with lactate oxidase and platinum blackwithin the anode and cathode, respectively, to achieve the powergenerating operation. Exemplary implementations of the exemplarywearable epidermal biofuel cell were performed to demonstrate theapplication of various forms of mechanical deformation relevant topractical epidermal applications, which resulted in minimal effects onthe performance of the device. For example, an exemplary implementationof the epidermal tattoo biofuel cell device during a controlled fitnessroutine revealed a maximum power density of 68 μW cm⁻² was obtained,hence realizing power production from human perspiration. The epidermalbioenergy paradigm thus holds noteworthy potential for use in thefitness, sport, athletics, performance, and generalized healthcaremonitoring domains.

As the cost of personal health monitoring continues to rise, the fitnessand healthcare industries have become increasingly reliant on wearablesensors to quantify various physiological metrics in a non-intrusive,user-friendly, and cost-effective fashion to reduce such costs. Forexample, for epidermal biosensing applications, durability,light-weight, and intimate skin conformance are core requirements ofsuch sensor devices to assess vital signs, e.g., such as heart rate,respiration rate, oxygenation of the blood, skin temperature, bodilymotion, brain activity, and blood pressure, as well as chemical sensorscapable of monitoring various physiological analytes on the wearer'sepidermis as well as chemical agents in their local vicinity. Forexample, these conformal electronic and diagnostic technologies haveadvanced considerably to the point of integration of disparate systemson a single skin-adhesive substrate. However, further progress in thisarena has been hindered by the lack of wearable and conformal powersources, especially those able to harness the mechanical or chemicalenergy produced by the wearer's body. While flexible and thin batterytechnologies have been developed, toxicity, longevity, device weight,and overall poor operational performance have precluded their use intransdermal applications, as well as the rigorous mechanical deformationencountered during bouts of physical activity remains to be addressedwith respect to these devices. Additionally, piezoelectric energyharvesting materials have also been plagued by the low efficienciesassociated with the electromechanical interconversion process incrystalline media lacking inversion symmetry. The disclosed wearableepidermal biofuel cell technology can be implemented to circumvent thesechallenges with conventional power sources and provide continuousextraction of biochemical fuels from the wearer's epidermis, which canfurther enable the development of epidermal electronics that can beutilized in the field.

Exemplary implementations of exemplary wearable epidermal biofuel celldevices were performed that demonstrated the ability to generate usefullevels of power from the perspiration of live subjects in a non-invasiveand continuous fashion through the use of temporary-transfer tattoos. Insome implementations, this was accomplished via the selective oxidationof lactate present in the wearer's perspiration through the inclusion ofthe enzyme lactate oxidase in the anode matrix in conjunction with thewater-insoluble electrochemical mediator tetrathiafulvalene (TTF). Forexample, lactic acid is the most abundant molecular constituent of theperspiration and is also a widely-recognized indicator of exerciseintensity, muscular exertion, fatigue, and aerobic/anaerobicrespiration. Charting lactate levels in real-time can thus yield timelyinformation regarding an individual's metabolic response to a fitnessroutine, hence enabling the individual, trainer, coach, and/orhealthcare provider to quantify performance levels. Advantageously, anindividual's fitness levels and aerobic capacity can indirectly beinferred by the amount of current (and hence power density) produced bythe device.

The disclosed tattoo biofuel cell devices address the requirementsimparted by epidermal wear, e.g., including, but not limited to, theability of the device to maintain its structural and electrochemicalresiliency against repeated (and often severe) mechanical deformationsuch as sheer stress and strain. For example, the exemplary tattoobiofuel cell devices can include dispersed carbon fibers within the inkused to print the anode and cathode electrodes, multi-walled carbonnanotubes incorporated in the electrode contingents to facilitateelectron transfer, as well as the immobilization of the catalyst (e.g.,lactate oxidase) entrapped in a biocompatible chitosan membrane, whichsynergistically results in the fabrication of biofuel cells that arelargely impervious to mechanical strain, stress, and degradationassociated with epidermal wear. For example, operation of the exemplarytattoo biofuel cell devices can produce a redox current from the directoxidation of lactate within the perspiration via biocatalysis at theanode (and concomitant catalytic reduction of oxygen at the cathode) togenerate electrical energy at a load. As such, the disclosed tattoobiofuel cell devices can be implemented in a number of practicalapplications to satisfy the energy requirements of epidermal,transdermal, and percutaneous devices.

Exemplary materials and methods to implement the disclosed embodiment ofthe technology are presented. The following chemicals and reagents wereused in the described implementations, which included tetrathiafulvalene(TTF), glutaraldehyde solution (8%), chitosan, Pt black, bovine serumalbumin (BSA), lactic acid, glucose, potassium phosphate monobasic(KH₂PO4), potassium phosphate dibasic (K₂HPO₄), hydrochloric acid (HCl),ammonium hydroxide (NH₄OH), sodium chloride (NaCl), potassium chloride(KCl), calcium chloride (CaCl₂), magnesium chloride (MgCl₂), and sodiumbicarbonate (NaHCO₃). Lactate oxidase (LOx) and carboxy-functionalizedmulti-walled carbon nanotubes (MWNTs-COOH) were obtained for use in theexemplary implementations. Exemplary reagents were used without furtherpurification. Carbon fibers (e.g., 8 μm diameter, 6.4 mm length, 93%purity) were obtained and further processed to reduce their length toapproximately 2 mm.

The fabrication of the exemplary tattoo biofuel cells used in thedescribed implementations included the following processes andprocedures, which were utilized in exemplary demonstrations andimplementations of the disclosed embodiment under exemplary conditionsdisclosed herein. Design of the temporary transfer tattoo pattern wasaccomplished in AutoCAD (Autodesk, San Rafael, Calif.) and fabricated on75 μm-thick stainless steel through-hole and mesh stencils (Metal EtchServices, San Marcos, Calif.). Unique stencil patterns were used foreach layer printed. Chopped carbon fibers were dispersed within aconductive carbon (E3449) ink to increase the tensile strength of theelectrode. Printing was performed using an MPM-SPM semi-automatic screenprinter (Speedline Technologies, Franklin, Mass.). Blank temporarytransfer tattoo paper and the accompanying adhesive substrate were used.

FIG. 46 shows a schematic illustration of an exemplary method tofabricate tattoo biofuel cells using screen printing techniques. Thefabrication method includes a process 4610 to deposit electrodes 4613 ona tattoo paper substrate comprising a release agent 4612 coated on abase paper 4611. For example, the release agent 4612 can includehydrophobic material that releases upon exposure moisture, e.g., such aspolydimethylsiloxane (PDMS), a cellulosic-based material, a siliconematerial, among others. For example, the electrodes 4613 can formed byscreen printing, roll-to-roll printing, aerosol deposition, inkjetprinting, or other printing techniques to fabricate a printed anode andcathode of the tattoo biofuel cell device. The electrodes 4613 can beformed of a carbon-based ink material or other electrically conductivematerial, which can include a catalyst, e.g., including, but not limitedto, an enzyme biocatalyst or noble metal catalyst, dispersed within theink. Implementation of the process 4610 to deposit the electrodes 4613can also include the formation of interconnects, contact pads, or otherelectrical components of the tattoo biofuel cell device. The process4610 can include a curing procedure to thermally or UV cure theelectrodes 1613 on the tattoo paper substrate material. In someimplementations, the process 4610 can include the deposition and curingof an underlayer of an electrically conductive material, which caninclude the interconnects, contact pads, or other electrical componentsof the tattoo biofuel cell device. The fabrication method includes aprocess 4620 to deposit a layer of a transparent insulator material 4624on the tattoo paper substrate exposing the electrodes 4613. Thefabrication method includes a process 4630 to modify the electrodes 4613with a biochemical modifier 4635. In some implementations, the process4630 can include attaching the catalyst as the biochemical modifier 4635to the anode and/or cathode by coating the catalyst as a layer on thesurface of the anode and/or cathode electrode; by entrapping thecatalyst in an electropolymerized conducting polymer formed on thesurface of the anode and/or cathode electrode; by entrapping thecatalyst using a selectively permeable scaffold-like structure, e.g.,such as an electro-permeable membrane, formed on the surface of theanode and/or cathode electrode; by covalently bonding the catalyst tothe surface of the anode and/or cathode electrode; or byelectrostatically anchoring the catalyst to the surface of the anodeand/or cathode electrode. In some implementations, the process 4630 caninclude attaching an electroactive mediator as the biochemical modifier4635, in addition to or alternatively to the catalyst, to the anodeand/or cathode electrode using any of the described techniques. Thefabrication method includes a process 4640 to deposit an adhesive layerof an adhesive material 4646 over at least a portion of the transparentinsulator material 4624 on the tattoo paper substrate, e.g., stillexposing the electrodes 4613, to produce the tattoo biofuel cell deviceready for implementation and wearable on a user's body. For example,subsequent to the fabrication method, the tattoo biofuel cell device canbe attached to a user in an on-skin transfer process 4650, in which theadhesive layer is directly attached to the skin and the tattoo papersubstrate is peeled off of the device by removing the release agent 4612(e.g., which also removes the base paper 4611). For example, in someimplementations, the fabricated tattoo biofuel cell device can include avoid region to permit sweat or other substance including the biofuel toflow.

The two electrode constituents of the tattoo biofuel cell were designedin the shape of ‘UC’ (acronym for the University of California). Asshown in FIG. 46, the entire contingent was printed on the tattoo basepaper using carbon fiber-reinforced (1.5% wt.) carbon ink via thethick-film screen printing fabrication process utilizing the stencilset. This was followed by the screen printing of a transparent insulator(Dupont 5036, Wilmington, Del.) on top of the carbon electrodes. Thestencil employed for the transparent insulator ink was designed toinsulate all but the active areas of the two electrodes. Following everyscreen printing step, the printed tattoo paper was cured at 90° C. for15 min in a convection oven.

Following the fabrication of the tattoo BFC, the anode (‘U’) wasmodified with LOx while the electrode ‘C’ was functionalized with Ptblack to serve as the cathode. With respect to the bioanodemodification, a suspension of carbon nanotubes in ethanol (5 mg/mL) wassonicated for several hours, and then mixed with 0.1 M TTFethanol/acetone solution in a 2.0:1.6 volume ratio. The suspension wassubsequently cast onto the open area of the anode. After the electrodescompletely desiccated, 5 μL LOx solution (40 mg/mL with 10 mg/mL BSA)was cast on the electrode, and then covered with 2 μL of 1 wt % chitosansolution. The electrodes were then cross-linked with glutaraldehydevapor and stored at 4° C. overnight. To modify the tattoo BFC cathode,an aqueous solution of 10 mg/mL Pt black was sonicated and 10 μL of thesuspension was cast on the electrode. Following complete desiccation, 1μL Nafion solution (5 wt %) was cast on the electrode to act as aprotective layer.

As illustrated in FIG. 46, in order to transfer the tattoos to asubstrate, a transparent adhesive sheet was first applied to the tattoopaper, which ensured that the tattoo adhered satisfactorily to thebody/substrate. A rectangular region was excised from the adhesive sheetsuch that the active anode and cathode areas remained unobstructed toenable the facile diffusion of lactate and oxygen to the respectiveelectrode contingents. In order to apply the adhesive layer to thesubstrate, one of the transparent protective sheets from the adhesivesheet was removed and the adhesive layer was first mated with plaintattoo base paper. Later, the second transparent protective sheet matedwith the adhesive sheet was removed to expose the adhesive layer. A voidwas also left between the anode and cathode contingents to facilitatethe flow of perspiration between these two components. Next, the tattoocontingent was applied to the substrate, the base paper was dabbed withwater to dissolve the release agent, and the wet base paper was gentlyremoved to expose the adhesive layer on the substrate. The tattoo BFCwas finally placed on the adhesive sheet already located on thesubstrate and removed by dabbing it with water and gently peeling thebase paper from the substrate.

In one exemplary embodiment of the disclosed tattoo biofuel cell device,an epidermal biofuel cell device includes a substrate formed of aflexible electrically insulative material structured to adhere to theskin of a user, an anode formed on the substrate of an electricallyconductive material, the anode including a catalyst to facilitate theconversion of a fuel substance in a biological fluid to a first productin an oxidative process that releases electrons captured at the anode,thereby extracting energy from the fuel substance, a cathode configuredon the substrate adjacent to the anode and separated from the anode by aspacing region, the cathode formed of a material that is electricallyconductive and capable of reducing an oxygenated substance in thebiological fluid to a second product in a chemical reduction process inwhich the second product gains electrons, and an anode electrodeinterface component and a cathode electrode interface component formedon the substrate and electrically coupled to the anode and the cathode,respectively, via electrical interconnects, in which the extractedenergy is addressable as electrical energy at the anode electrodeinterface component and the cathode electrode interface component.

Exemplary implementations of the exemplary tattoo biofuel cell devicewere performed to perfect the device with regards to the electrochemicalperformance in vitro. For example, the tattoo biofuel cells were firstevaluated by transferring the pattern onto a rigid plastic substrate oronto a flexible GORE-TEX textile for mechanical integrity studies. Forexample, 0.2 M McIlvaine buffer (pH 5.5) was utilized to emulate theaverage pH value of human perspiration. With respect to in vitrostability evaluation, artificial perspiration was prepared with thefollowing electrolytes, metabolites, and small molecules, e.g.,including Na₂SO₄, NaHCO₃, KCl, MgCl₂, NaH₂PO₄, CaCl₂, acetic acid,lactic acid, pyruvic acid, glucose, uric acid, urea, creatinine andascorbic acid. The pH of the artificial perspiration stock solution wasadjusted to 5.3 by 5 M NH₄OH. The exemplary solutions were prepared withultra-pure water (18.2 MΩ·cm). Electrochemical characterization wasperformed at room temperature leveraging a CH Instruments (Austin, Tex.)model 1232A potentiostat.

Healthy volunteer subjects participated in the exemplary powergeneration experiments. Each volunteer was instructed to wear atemporary transfer tattoo BFC on their upper bicep in order to assessreal-time power generation. The BFC was connected to an external 100 kΩload resistor (R_(L)) in order to achieve maximum power transfer. Thisvalue was selected to most closely match the internal series resistance(R_(b)) such that the maximum power transfer condition was satisfied(R_(s)=R_(L)). Electrical current was recorded every 5 s using aKeithley (Cleveland, Ohio) 6514 system electrometer interfaced with acomputer system including at least a processor and a memory unitincluding a control program (e.g., instructions in Matlab) tocontinuously process acquired current readings via the GPM interface andinterpolated the concomitant power generated per unit area(P_(DENSITY)=I²R_(L)/A_(E), A_(E)=0.06 cm²). In order to filterextraneous noise, a 10-point moving average was iterated at each datapoint. For example, the subjects were instructed to mount a stationarycycle and a heart rate (HR) monitor was employed to track the subjects'HR. Subjects were instructed to begin cycling at a steady, slow cadencefor 3 min. Following this ‘warm-up’ period, subjects were instructed tocycle at an increasing pace until 80% of their maximum heart rate wasachieved in order to ensure that the anaerobic respiration threshold wasattained, hence augmenting the excretion of lactic acid in theperspiration. Immediately following the subjects' transition to theanaerobic regime, the subjects were instructed to maintain their currentcadence for 15 min in order to observe the temporal evolution of thelactate level. Following the 15 min intense exercise activity, subjectswere instructed to gradually reduce their cadence during a 3 min‘cool-down’ period. The volunteers ingested no fluid (dehydrated state)prior to and during the duration of the fitness routine.

Exemplary implementations of exemplary wearable epidermal biofuel celldevices were performed that demonstrated resiliency against mechanicalstress caused by continuous body movements. For example, the longevityof such epidermal-mounted devices depend greatly on their ability toadhere well to the human skin without developing fractures that damagethe devices. The most common body movements involve flexions, whichtypically comprise of bending, stretching, and/or twisting of theepidermal layer. Accordingly, such devices must encompass an intrinsicflexible and stretchable nature in addition to being able to adhere wellto the epidermis. The disclosed tattoo biofuel cell devices includedispersion of carbon fibers within the inks employed to print thesedevices, which provide a conductive, interleaved backbone that aids inmaintaining the electrical conductivity under various biomechanicalstressors. Similarly, the use of an adhesive layer firmly attaches thetattoo biofuel cells to the skin. Visual analysis of the tattoo biofuelcell device on the dorsal region of a human wrist under repeatingbending, stretching, and twisting dorsiflexion movements was performedfor a total of 50 iterations. FIG. 47 shows images of the epidermaltattoo biofuel cells during mechanical stress caused by continuous bodymovements including (i) stretching, (ii) bending, and (iii) twisting.The left column of image in FIG. 47 provides the side view of dorsalmovements; the middle column provides images of the top view of thebiofuel cell tattoos during the various deformations; and the rightcolumn provides images of the top view of the biofuel cell tattoos atthe end of each movement. The images demonstrate that the tattoo biofuelcell devices are quite resilient to flexions that emulate epidermalwear, e.g., as a consequence of the overlying insulator layer, whichserves to maintain the structural integrity of the printed carbon layer.Accordingly, the epidermal biofuel cell devices can perform desirablyunder various strains and thus can serve as a compelling platform forvarious epidermal applications.

To date, the majority of lactate biofuel cells have been based on thelactate dehydrogenase enzyme. However, in these existing devices, NAD⁺must be employed as the cofactor, which represents a noteworthychallenge given that this molecule must be immobilized on the electrodeto prevent it from leeching into the matrix while being able to diffuse,with relative ease, to the enzyme's active site.

The disclosed technology includes an exemplary lactate-based biofuelcell utilizing the lactate oxidase (LOx) enzyme for non-invasive powergeneration from human perspiration, e.g., by selectively catalyzing theoxidation of lactate in the perspiration as the biofuel for epidermalpower generation. In some implementations the electrodes of theexemplary tattoo biofuel cell device are functionalized to achieveefficient bioelectrocatalytic conversion, e.g., in which the ‘U’ of thetattoo (anode) was functionalized with MWNTs/TTF/LOx, hence serving asthe bioanode to catalyze the oxidation of lactate to pyruvate in thepresence of oxygen (cofactor). The cathode ‘C’ made use of a drop-castedPt black layer, protected with a Nafion proton-exchange membrane.

An image of the exemplary functionalized device is shown in FIG. 48A.The bioanode of the exemplary tattoo biofuel cell device wasfunctionalized with the mediator TTF and MWNTs. This bioanode was thencovered with a layer of chitosan, e.g., a naturally-derived biopolymerwell-known for its biocompatibility. For example, chitosan not onlyserves to protect the modified enzyme electrode, but it also functionsas a physical barrier to limit the efflux of the biocatalytic backbonefrom the tattoo and onto the underlying substrate.

FIG. 48B shows a data plot of polarization curves of the exemplaryfunctionalized MWNTs/TTF/LOx bioanode in the absence of presence of 14mM lactic acid in 0.2 M McIlvaine buffer solution, pH 5.5, respectively.The electrocatalytic activity of the MWNTs/TTF/LOx bioanode wasdetermined in vitro with an external Ag/AgCl (1 M KCl) electrode and aPt wire counter electrode. Polarization curves were recorded by applyinglinear sweep voltammetry with a scan rate of 1 mV/s in McIlvaine bufferpH 5.5 with 14 mM lactic acid, and normalized by the surface area of theelectrode as a function of potential. As shown in FIG. 48B, theTTF-mediated oxidation of lactic acid initiates from around −0.1 V witha peak potential of 0.14 V (vs. Ag/AgCl), indicating that theMWNTs/TTF/LOx exhibits selective catalytic ability towards the oxidationof lactic acid, and hence serves as a suitable bioelectrocatalyticcascade for the bioanode constituent of the BFC. For example, TTF can beused a selective mediator to aid in electron transfer between the LOxactive site and the electrode surface. Other mediators, e.g., including,but not limited to, derivatives of ferrocene and Meldola's blue, canalso be used as the selective mediator of the bioanode. It is noted thatalthough able to mediate the electro-oxidation of lactic acid, theseother small-molecule mediators are water-soluble, and the oxidationcurrent obtained may be decayed as a consequence of the leaching of themediator. Compared with these mediators, TTF encompasses severalnoteworthy advantages, namely lower oxidation potential and more stableperformance. Also, the incorporation of MWNTs further shifted the lacticacid oxidation onset potential more negatively and further enhanced theoxidation current, which may be due to the electron donor-acceptorinteraction between TTF and negatively charged MWNTs, resulting infacilitated electron transfer to the electrode. Therefore, theMWNTs/TTF/LOx cascade is well-suited to serve as the bioanode, and,together with a Pt black cathode, a complete lactic acid biofuel cellcan be assembled on the exemplary temporary transfer tattoo substrate.

FIG. 48C shows a data plot of power density achieved from the exemplarytattoo biofuel cell device with different lactic acid concentrations. Asshown in the figure, the exemplary tattoo biofuel cell device approached25 μW cm⁻² with 8 mM lactic acid (dissolved in buffer), and increased to34 and 44 μW cm⁻² with further increased lactic acid concentrations of14 mM and 20 mM, respectively. A small signal was observed duringcontrol experiments (no lactic acid added).

While this patent document contain many specifics, these should not beconstrued as limitations on the scope of any invention or of what may beclaimed, but rather as descriptions of features that may be specific toparticular embodiments of particular inventions. Certain features thatare described in this patent document in the context of separateembodiments can also be implemented in combination in a singleembodiment. Conversely, various features that are described in thecontext of a single embodiment can also be implemented in multipleembodiments separately or in any suitable subcombination. Moreover,although features may be described above as acting in certaincombinations and even initially claimed as such, one or more featuresfrom a claimed combination can in some cases be excised from thecombination, and the claimed combination may be directed to asubcombination or variation of a subcombination.

Similarly, while operations are depicted in the drawings in a particularorder, this should not be understood as requiring that such operationsbe performed in the particular order shown or in sequential order, orthat all illustrated operations be performed, to achieve desirableresults. Moreover, the separation of various system components in theembodiments described in this patent document should not be understoodas requiring such separation in all embodiments.

Only a few implementations and examples are described and otherimplementations, enhancements and variations can be made based on whatis described and illustrated in this patent document.

1.-20. (canceled)
 21. A method of sensing an analyte in a biofluid usinga non-invasive epidermal electrochemical sensor device, comprising:passing an electrical signal between a first iontophoretic electrode inan anode electrode assembly of the non-invasive epidermalelectrochemical sensor device and a second iontophoretic electrode in acathode electrode assembly of the non-invasive epidermal electrochemicalsensor device to create an electric field to drive ion flow from sweatpresent on skin of a user of the non-invasive epidermal electrochemicalsensor device toward a working electrode and at least one of a counterelectrode or a reference electrode in the anode electrode assembly ofthe non-invasive epidermal electrochemical sensor device; catalyzing ananalyte in the sweat via a catalyst in an electrochemical transducerlayer on the working electrode in the anode electrode assembly of thenon-invasive epidermal electrochemical sensor device to cause a reactiondetectable at the anodic and cathodic electrode assemblies, andtransmitting, via an electrically conductive contacts in an electrodeinterface assembly coupled to the electrodes of the anodic electrodeassembly and the cathodic electrode assembly, a sensor signal indicativeof the reaction to one or more electrical circuitry.
 22. The method ofclaim 21, including: detecting the analyte in the sweat based on thesensor signal at the one or more electrical circuitry.
 23. The method ofclaim 22, wherein the one or more electrical circuitry is integratedwith the non-invasive epidermal electrochemical sensor device, and themethod further comprises: wirelessly transmitting, from the non-invasiveepidermal electrochemical sensor device, data associated with thedetected analyte to an external device.
 24. The method of claim 21,wherein the passing the electrical signal between the first and secondiontophoretic electrodes to create the electric field includes causing asweat-inducing chemical agent to be driven from a hydrogel or cryogellayer covering at least one or both of the first and secondiontophoretic electrodes onto a target skin area.
 25. The method ofclaim 21, wherein the sensor signal indicative of the reaction isgenerated using one of an amperometric technique, a potentiometrictechnique, a conductometric technique, or a voltammetric technique. 26.A method of inducing sweat and sensing an analyte in the sweat using anon-invasive epidermal electrochemical sensor device, the methodincluding: passing an electrical signal between a first iontophoreticelectrode in an anode electrode assembly of the non-invasive epidermalelectrochemical sensor device and a second iontophoretic electrode in acathode electrode assembly of the non-invasive epidermal electrochemicalsensor device to create an electric field to administer a sweat-inducingchemical gent from the anode electrode into skin of a user of thenon-invasive epidermal electrochemical sensor device to cause localizedsweat generation near a working electrode and at least one of a counterelectrode or a reference electrode in the anode electrode assembly ofthe non-invasive epidermal electrochemical sensor device, wherein theelectric field drives ion flow from the generated sweat excreted on theskin; catalyzing an analyte in the sweat via a catalyst in anelectrochemical transducer layer on the working electrode in the anodeelectrode assembly of the non-invasive epidermal electrochemical sensordevice to cause a reaction detectable at the working electrode, andtransmitting, via an electrically conductive contacts in an electrodeinterface assembly coupled to the electrodes of the anodic electrodeassembly and the cathodic electrode assembly, a sensor signal indicativeof the reaction to one or more electrical circuitry.
 27. The method ofclaim 26, including: detecting the analyte in the sweat based on thesensor signal at the one or more electrical circuitry.
 28. The method ofclaim 27, wherein the one or more electrical circuitry is integratedwith the non-invasive epidermal electrochemical sensor device, and themethod further comprises: wirelessly transmitting, from the non-invasiveepidermal electrochemical sensor device, data associated with thedetected analyte to an external device.
 29. The method of claim 26,wherein the passing the electrical signal between the first and secondiontophoretic electrodes to create the electric field includes causingthe sweat-inducing chemical gent to be driven from a hydrogel or cryogellayer covering at least one or both of the first and secondiontophoretic electrodes onto a target skin area, and wherein thesweat-inducing chemical agent includes pilocarpine.
 30. The method ofclaim 26, wherein the sensor signal indicative of the reaction isgenerated using one of an amperometric technique, a potentiometrictechnique, conductometric technique, or a voltammetric technique. 31.The method of claim 21, wherein the non-invasive epidermalelectrochemical sensor device includes an adhesive membrane and aflexible or stretchable substrate disposed over the adhesive membrane,wherein the anodic electrode assembly and the cathode electrode assemblyare disposed over the flexible or stretchable substrate, and wherein theadhesive membrane includes a gel layer and at least one of a siliconemembrane, a urethane membrane, or an acrylic adhesive membrane.
 32. Themethod of claim 31, wherein the gel layer includes a charged chemicalagent, and the charged chemical agent includes a sweat-inducing chemicalagent.
 33. The method of claim 21, wherein the detected analyte includesascorbic acid.
 34. The method of claim 21, wherein the detected analyteincludes glucose.
 35. The method of claim 21, wherein the detectedanalyte includes creatine.
 36. The method of claim 26, wherein thenon-invasive epidermal electrochemical sensor device includes anadhesive membrane and a flexible or stretchable substrate disposed overthe adhesive membrane, wherein the anodic electrode assembly and thecathode electrode assembly are disposed over the flexible or stretchablesubstrate, and wherein the adhesive membrane includes a gel layer and atleast one of a silicone membrane, a urethane membrane, or an acrylicadhesive membrane.
 37. The method of claim 36, wherein the gel layerincludes a charged chemical agent, and the charged chemical agentincludes a sweat-inducing chemical agent.
 38. The method of claim 26,wherein the detected analyte includes ascorbic acid.
 39. The method ofclaim 26, wherein the detected analyte includes glucose.
 40. The methodof claim 26, wherein the detected analyte includes creatine.